Combinational multiphoton scanning microscopy and multiphoton surgery of mouse arteries. Samira Karimelahi

Transcription

1 Combinational multiphoton scanning microscopy and multiphoton surgery of mouse arteries by Samira Karimelahi A thesis submitted in conformity with the requirements for the degree of Masters of Applied Sciences Graduate Department of Electrical and Computer Engineering University of Toronto Copyright c 2011 by Samira Karimelahi

2 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 laserstimulated 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 lasersemitransparent materials interaction were performed. A purpose-built two-photon fluorescence 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

3 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. iii

4 Contents 1 Introduction Thesis objectives Chapter by chapter outline Background Laser interaction with transparent materials Nonlinear ionization Femtosecond laser modification inside transparent materials Cell and tissue disruption by femtosecond laser Applications Two-photon fluorescence imaging Mechanism of multiphoton fluorescence microscopy Architecture of two-photon fluorescence microscope Multiphoton imaging applications Experiment Femtosecond laser system Beam delivery system Purpose-built two-photon fluorescence imaging setup Hardware of the fluorescence microscope system TCSPC laser microscope software iv

5 3.4 Sample preparation Mouse blood vessel loaded in the microfluidic chip Fluorescent microspheres Results and discussion Two photon fluorescence imaging of microspheres TPI by two dry lenses with different numerical apertures TPI by oil-immersion lens Two photon fluorescence imaging of the waveguides inside the fused silica Two photon fluorescence imaging of the optical fiber Breakdown threshold at various microfluidic chip interfaces Damage threshold: glass-air interface Damage threshold: glass-pdms interface Damage threshold: glass-mops interface Comparison between damage threshold results for different interfaces Bubble formation threshold inside MOPS solution Micromachining on the mouse artery wall Combination of trepanning and two photon fluorescence imaging Conclusion Summary Future directions A Light propagation in the matter 112 B Two photon absorption probability 115 Bibliography 116 v

6 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

7 List of Tables 3.1 Optical focusing parameters related to three lenses Comparison of calculated beam spot size and depth of focus as well as the measured lateral and axial length of the fluorescent beads Comparison of calculated lateral and axial resolution according to the Rayleigh Resolution criteria definition Laser exposure parameters for writing waveguides inside fused silica Threshold pulse energy and irradiance for different microfluidic chip interfaces in single and a multiple-pulse laser exposure vii

8 List of Figures 2.1 Schematic models of different kinds of the photoionization according to Keldysh parameter, γ [31] 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 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] 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] Diagram of important parameters in laser-tissue interaction [40] Absorption spectra for the three dominant components: water, hemoglobin and melanin in tissue [41] viii

9 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] (a) Ablation lines with five different pulse energies on fluorescently-labeled actin fibers. (b) Fluorescence intensity profile with respect to position along sample [50] 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 mm 2 with depth of 200 µm, and the second cut removed an area of 0.25 mm 2 with a final depth of 360 µm [51] Two-photon fluorescence images of the brain s blood vessel disruption using 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] 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] (a) Wavelength distribution of fluorescence and second harmonic signal (b) isotropic emission of the fluorescence signal [12] Dependence of the excitation process on axial distance for one photon and two-photon excitation [11] Two-photon fluorescence imaging set up [12] ix

10 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] Multiphoton imaging of the vascular distribution of (a) normal and (b) tumor blood vessels [74] Temporal evolution of drug delivery technique imaged by two-photon fluorescence imaging [75] 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] Fiber chirped-pulse amplification arrangement of the fiber fs laser [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 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] Npaq Control Assembly showing various jumpers. The jumper JP1 was changed from the default position of 1-2 to Depth correction of refraction of the light according to the Snell s law at the interface of two materials with different refractive indices n 1 and n Two-photon fluorescence imaging setup. See text for detailed explanations of components x

11 3.7 Hardware block diagram of the imaging system (modified from [86]) Principle of the TCSPC measurement [88] Three trigger pulses that determine pixel, line, and frame [86] Control and Analysis tab showing different features of the TCSPC Laser Microscope Software [87] (a) 3D display of the sample two-photon fluorescence image. (b) 3D intensity profile [87] Image acquisition software view showing a 2D image of multiple 1 µm microspheres Schematic illustration of the different steps of fabricating PDMS stamp. Modified from [93] 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] Images of 1 µm diameter fluorescent beads (a) under SEM, and (b) under a bright field microscope with 100 objective lens Cross section TPI of the 1 µm fluorescent beads recorded with 40X-0.65 NA The axial intensity profile of one of a fluorescent bead represent in (a) ImageJ and in (b) MATLAB (red line) Scanning laser beam through the microsphere beads mounted on the microscope slide The x-y image of the fluorescent beads recorded with a 0.65 NA lens: each pixel is equivalent to 0.5 µm size 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

12 4.6 The lateral intensity profile of a fluorescence sphere (1 µm diameter) observed and represented by a Gaussian curve (red line) with MATLAB tools 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 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 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 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 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 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] Single mode fiber optic taped on a microscope slide for TP laser image TPI of a optical fiber with a 40X-0.65 NA dry lens (a) cross section (xz) and (b) top view (xy) 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) Optical microscopic images of femtosecond laser line scan on the top surface of the cover slip interface with air. Each average power line scan is repeated for scan speeds from 0.2 to 50 mm/s xii

13 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 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 Minimum number of pulses corresponding to the specific pulse energy required 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 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 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 Comparison between laser damage threshold on the bottom and top surfaces of the cover slip for glass-to-air interfaces 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] 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

14 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 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 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 Comparison between damage threshold of different interfaces at microfluidic chip for various exposure condition 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] The minimum number of laser pulses inducing bubble formation inside the MOPS solution versus average power. The solid line is a guide Comparison of all Breakdown thresholds at different exposure conditions Mouse artery wall loaded in the microfluidic chip. The laser beam ablated the blood vessel in the x direction 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 Scanning laser beam in the z direction (vertical to blood vessel surface) on the artery wall µm hole created by laser power of (a) 80 mw laser beam and (b) 70 mw.103 xiv

15 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 Cross sectional two photon fluorescence image of the unstained artery wall. Scale is the same for both directions 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 Iodide. b) Cross section (x-z) image of the blood vessel with both Propidium Iodide (right image) and Fura-Red stained (left image) The CCD camera images of the artery wall (a) before and (b) after laser exposure Two photon fluorescence imaging of the artery wall after laser trepanning exposure. The diameter of the hole is around 80 µm xv

16 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 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 attention 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

17 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 sample. 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 fluorescence 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 nonlinear excitation only in the focal volume, three dimensional image sectioning is obtainable 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 applications 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 activity [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 exposure 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

18 Chapter 1. Introduction 3 properties make microfluidic chips appropriate for analyzing single cells, cellular structures, 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 transparent 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 objective 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

19 Chapter 1. Introduction 4 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 different interfaces as well as the bubble formation threshold in the MOPS (physiologic 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.

20 Chapter 1. Introduction 5 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 focusing on the nonlinear processes induced by the femtosecond laser. The ultrashort pulses can be used to modify sample structures in both semiconductors and biological 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 delivery path. Also, our purpose-built fluorescence imaging setup is explained in detail for both hardware and software aspects. Moreover, the sample preparation methods 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.

21 Chapter 1. Introduction 6 Chapter 5 Conclusion presents a summary of this research work and its significance as well as possible future directions.

22 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

23 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 absorption 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] 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-

24 Chapter 2. Background 9 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 me cnɛ 0 E g, (2.1) e I where ω is the laser frequency, m e 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, ɛ 0 is the permittivity of free space, and E g 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:

25 Chapter 2. Background 10 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 ionization 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 electrons 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] Femtosecond laser modification inside transparent materials 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

26 Chapter 2. Background 11 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 I N in the material with the band gap (Eg) equivalent to N photons energy satisfying Nhν = E s. 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 parameters that affect the damage threshold intensity [36]. One of the complications in measuring damage threshold inside the transparent material 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 increase, until it reaches the critical power (Eq. (2.3)) in which self focusing balances the

27 Chapter 2. Background 12 diffraction and creates a filament. The critical power is given by: P cr = 3.77λ2 8πn 0 n 2, (2.3) where λ is a laser wavelength, and n 0 and n 2 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 created 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 properties like absorption and scattering coefficient, heat capacity, and thermal conductivity. Laser parameters like energy pulse, repetition rate, wavelength, and beam spot size will

28 Chapter 2. Background 13 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].

29 Chapter 2. Background 14 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

30 Chapter 2. Background 15 Figure 2.5: Diagram of important parameters in laser-tissue interaction [40].

31 Chapter 2. Background 16 Figure 2.6: Absorption spectra for the three dominant components: water, hemoglobin and melanin in tissue [41].

32 Chapter 2. Background 17 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

33 Chapter 2. Background 18 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]. Ultrashort 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 temperature 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 threshold, 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

34 Chapter 2. Background 19 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 transparent 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. Therefore, 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 nonequilibrium 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 cm 3 [48]. So, it is necessary to understand 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.

35 Chapter 2. Background 20 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 µm [50]. s 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 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.

36 Chapter 2. Background 21 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] 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 disturbances can happen as is described in Fig [52]. One disturbance is blood plasma

37 Chapter 2. Background 22 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 mm 2 with depth of 200 µm, and the second cut removed an area of 0.25 mm 2 with a final depth of 360 µm [51].

38 Chapter 2. Background 23 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 using 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

39 Chapter 2. Background 24 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 theoretically 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 twophoton 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] 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]. Twophoton 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

40 Chapter 2. Background 25 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

41 Chapter 2. Background 26 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 wavelength 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 second 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].

42 Chapter 2. Background 27 As it is shown in detail in Appendix B, one important factor in two-photon absorption probability, n a, is the two-photon cross section which is different for each fluorophore. A larger two-photon cross section, σ 2, results in higher two-photon absorption rate, so it is important to select the fluorophore with large σ 2 that can be excited by the available laser source [61]. Multiphoton absorption depends nonlinearly on the intensity. This intensity dependence makes multiphoton absorption localized. As shown in Eq. (B.4), two-photon absorption is proportional to the square of the intensity. This quadratic dependence originates 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 appropriate detector can collect the fluorescence signal coming from the interaction volume in the sample [11] Architecture of two-photon fluorescence microscope An arrangement for two-photon fluorescence imaging is shown in Fig One important 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

43 Chapter 2. Background 28 Figure 2.13: Dependence of the excitation process on axial distance for one photon and two-photon excitation [11].

44 Chapter 2. Background 29 femtosecond pulse duration with moderate average power. Both picosecond and femtosecond 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 elements in microscopy which determines the speed of the capturing frame. Computercontrolled 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 parameters are spectral range, electronic noise level, cost, readout speed, and quantum efficiency. Photomultiplier tubes (PMT), avalanche photodiode (APD), and charge-coupled detectors 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

45 Chapter 2. Background 30 front of the detector to achieve optical sectioning for constructing three dimensional images. 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] 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 dynamics 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 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

46 Chapter 2. Background 31 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].

47 Chapter 2. Background 32 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 distribution [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 disruption 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 vasculature, mice were injected intravenously with fluorescent dyes. Fig 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 micronanosurgery 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

48 Chapter 2. Background 33 Figure 2.17: Temporal evolution of drug delivery technique imaged by two-photon fluorescence 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 [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].

49 Chapter 2. Background 34 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 biology 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 fabricate 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-

50 Chapter 2. Background 35 ance for in-situ laser trimming or post-processing as well as real-time feedback for controlling laser fabrication process [81].

51 Chapter 3 Experiment Experiments carried out for this work include two-photon fluorescence imaging and femtosecond 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, femtosecond 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 purposebuilt 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

52 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 khz Ti:Sapphire regenerative amplifiers and low pulse energy 80 MHz Ti:Sapphire oscillators [2]. The fiber chirped-pulse amplification technology used in this laser is shown in Fig 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 M 2, 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 (z R ). Subsequently, the beam divergence half-angle can be calculated via λ πz R. For our system, M 2 was calculated by my colleagues to be 1.3.

53 Chapter 3. Experiment 38 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 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 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.

54 Chapter 3. Experiment 39 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

55 Chapter 3. Experiment 40 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 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].

56 Chapter 3. Experiment 41 Figure 3.4: Npaq Control Assembly showing various jumpers. The jumper JP1 was changed from the default position of 1-2 to 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 advantage 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 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-

57 Chapter 3. Experiment 42 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: NA dry lens (Nikon, BD plan 100X/0.9) and 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, n 1 and n 2 that are refractive indices of the two mediums is given by Eq. (3.5). d 1 = n 1 ( NA 1 n 1 ) 2 (3.1) d 2 n 2 1 ( NA n 2 ) 2 where d 1 and d 2 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 M 2 π NA. (3.2) Here, we have assumed the beam coming to the lens is collimated, and NA is the

58 Chapter 3. Experiment 43 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 n 1 and n 2. numerical aperture of the lens, λ is the laser wavelength, and M 2 is the beam quality factor. The Rayleigh range (z R ) of the Gaussian beam with the diffraction-limited assumption can be defined as: z R = π w2 n. (3.3) λ 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 d 1 /d 2, the beam spot size, Rayleigh range, and the depth of focus are summarized in Table 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 coupled 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

59 Chapter 3. Experiment 44 Lens specifications d 1 /d 2 Beam spot size [µm] Rayleigh range [µm] Depth of focus [µm] Nikon, 40X 0.65 NA Nikon, 100X 0.9 NA Nikon, 100X 1.25 NA 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 ), 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 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 The optical

60 Chapter 3. Experiment 45 Figure 3.6: Two-photon fluorescence imaging setup. See text for detailed explanations of components. path in this block diagram refers to the beam delivery system that guided the laser pulses from the laser source to the sample and the imaging setup which collected the fluorescence signal and sent it to the detector via a single mode fiber. The detected signal which was converted to the electronic signal, a trigger timing signal from the laser, and trigger reference positions from the motion stage were sent to the computer including the time correlated single photon counting (TCSPC) board. A synchronization TTL output signal from the laser head was used as a reference clock to synchronize the laser induced fluorescence signal while three trigger signals from the motion stage were related to the capturing time of the pixel, line, and frame. The TCSPC board created the image of the sample using these inputs. More details of the TCSPC method will be discussed in the following sections.

61 Chapter 3. Experiment 46 Figure 3.7: Hardware block diagram of the imaging system (modified from [86]).

62 Chapter 3. Experiment 47 Time-Correlated Single Photon Counting (TCSPC) For detecting the fluorescence signal, the Time-Correlated Single Photon Counting (TC- SPC) technique has been applied. The system has been designed to acquire photon distributions versus spatial coordinates of the XYZ stages. The trigger of capturing pixel, line and frame was sent from the motion stage to the TCSPC board. Also, the comparator circuit was implemented in order to convert the differential signals sent by the motion stage controller to TTL signals as a trigger input. TCSPC board was used to detect fluorescent photons through the Avalanche Photodiode detectors (APD) [86, 87]. As illustrated in Fig. 3.8, the TCSPC technique is based on the number of photons detected in the specific time interval. Each time interval consists of small time periods in which the probability of multiple photon detection can be neglected for low level signal there the one photon detection probability is much less than one. The signal coming to the detector consists of random pulses related to the detected fluorescence photons. In some time periods no photon is detected. Every time that a photon is detected, a number is saved in the memory address proportional to the detection time. Consequently, the number of photons detected in each time interval is recorded. The histogram of the detection times and the number of photons in each time interval will reconstruct the waveform. As described here, the fluorescence signal coming from the fluorophore could be detected using this method, and the intensity level of each pixel could be constructed accordingly to produce the time dependent waveform of the fluorescence emission [88]. In our system, a multi-dimensional TCSPC system was used which could record photon distribution of up to five dimensions that are lifetime, wavelength, and the X, Y, and Z coordinates of the scanning area. Three triggering pulses are shown in Fig. 3.9 [86, 87]. These three triggers define the pixel, line, and frame values that were sent from the motion controller to TCSPC system in order to get the photon distribution versus the spatial coordinates in the target sample.

63 Chapter 3. Experiment 48 Figure 3.8: Principle of the TCSPC measurement [88]. Figure 3.9: Three trigger pulses that determine pixel, line, and frame [86].

64 Chapter 3. Experiment TCSPC laser microscope software In order to create an image from the laser raster scanning, two software programs were used. The first software was the motion controller software (Aerotech A3200, Nview MMI) which controls the stage and scans the sample via G-code. The second software was image acquisition software (Becker & Hickl, SPC-830) that acquired the scanned fluorescence data via the photon counter, and displayed and analyzed the acquired 2D images. Image acquisition software software was available on the same computer as the TCSPC board from the Becker & Hickl Company, was installed while the motion controller software was available on an other computer to control the motion stages. By networking the computer containing the image acquisition software and computer containing motion control software via a LAN connection, one could use just one software with a Labview interface to control the motion stage, acquire the image via the photon counter, analyze the acquired 2D images, and use volume rendering to plot multiple 2D slices into one 3D model. The image acquisition computer was chosen to be the Master computer, whereas the motion stage computer was set as the Slave. The user was able to control the whole system via the Master computer alone. As a result, we had a single, efficient, powerful, and user friendly software module for our TCSPC Laser microscope system. A view of the control and image analysis tab on the Labview interface software is shown in Fig The software has three different tabs: initializing, which puts the motion stage at a reference position; and control and image analysis that is the main feature of this software which includes analysis of the image to provide intensity plots as well as setting the scan speed, pixel size, and other data collection options. By clicking on one of the icons of default raster scan, scan from file or costume scan, the related G-code from the Master computer was sent to the Slave one. The slave computer compiled the G-code file and sent it to the motion stage. Perfect time synchronization was achieved using pixel, line and frame trigger signals that were sent from

65 Chapter 3. Experiment 50 Figure 3.10: Control and Analysis tab showing different features of the TCSPC Laser Microscope Software [87].

66 Chapter 3. Experiment 51 the motion stage to the photon counter on each move. The photon counter image acquisition software (SPC-830) provided functions such as measurement control, calculation and display of data, set-up of measurement parameters, 2-dimensional and 3-dimensional display of measurement results, mathematical operations, selection of subsets from multidimensional data sets, loading and saving of results and system parameters, and control of the measurement in the selected operation mode. Some of the features such as 3D display of the image and 3D intensity profile are shown in Fig Figure 3.11: (a) 3D display of the sample two-photon fluorescence image. (b) 3D intensity profile [87]. There were some defects in the joint software program with the Labview interface that prevented us to record images. Therefore, the images in this thesis were collected by running the motion controller and image acquisition softwares independently, but concurrently in order to capture the image. The connection problem has been solved recently but was not available for this thesis work. Two slightly different G-codes have been used for imaging the cross section (x-z) and the top view (x-y). In each code, the start position which is the center of the image was set by the user. The default settings of the program were set to acquire a frame

67 Chapter 3. Experiment 52 of pixels. The size of the pixel also was set manually. To take the top view image, the step size for both direction (X and Y) is the same because the focus position is changing inside one material. However, for the cross section image, the step size in the vertical direction was different from the horizontal one. The correction on the step size carried out according to Eq. (3.5). In order to capture a 2D image in the x-z or x-y plane, the stage was moved 128 X step size in the X direction, and 128 Z step size or Y step size in Z or Y direction from the starting position. From this position which is the pixel 1 1, the software starts to acquire data and measure the number of detected photons per pixel. This process will end by capturing data for the pixel At the end, a 2D display of the image will be created by the image acquisition software. An example of the 2D image of the sample presented by image this acquisition software is shown in Fig Sample preparation Mouse blood vessel loaded in the microfluidic chip Microfluidic devices are useful in handling small samples, and integrating multiple processes required for lab-on-a-chip experiments. These properties make them appropriate for analyzing single cells and cellular structures. Several research areas spanning from biological analysis and sensing, to optics and photonics, can be influenced by the developments in microfluidic devices [89,90]. Furthermore, the geometries offered by microfluidic chips result in spatial confinement, as well as the creation of biochemical gradients that better mimic the in vivo environment [91]. Due to the ease of processing, the ability to support certain useful components (such as pneumatic valves) and low material cost, most of the research to date has focused on fabricating such devices in PDMS substrates, which are an optically transparent and soft elastomer [89, 92]. The microfluidic chip designed by Professor Axel Guenther group, our collaborator

68 Chapter 3. Experiment 53 Figure 3.12: Image acquisition software view showing a 2D image of multiple 1 µm microspheres.

69 Chapter 3. Experiment 54 from Mechanical Engineering Department, has several unique features: (1) holding the blood vessel segment directly against the glass cover where ideal focusing and precise nanosurgery is possible and (2) matching the channel size to the blood vessel, allowing its immobilization for precise probing. One of the most powerful patterning techniques for microfabrication is photolithography. As the entire pattern of the photomask can be copied onto a thin film of photoresist at the same time, projection photolithography is a parallel process. Mass-production is one of the features of photolithographic techniques with the capability of projection of patterned structures of about 100 nm size in thin films of photoresists. Because photolithography is not inexpensive, directly applicable to all photosensitive materials, and suitable for three-dimensional structure, it is not the best option for all applications. Soft lithography, on the other hand, uses patterned elastomer as the stamp, mold, or mask to generate micropatterns instead of the rigid photomask, which makes it a suitable substitute for photolithography. Moreover, this method is capable of creating structures with a few nanometric resolution. Being low cost, straightforward, capable of qausi-three dimensional structures, and applicable for a wide variety of materials makes the soft lithography attractive as a flexible fabrication tool [93]. A microfluidic chip was fabricated by Professor Axel Guenther from mechanical engineering department. The microfluidic devices were designed and fabricated using single and multilayer soft lithography in the elastomeric substrate material polydimethylsiloxane (PDMS). Their method for making microfluidic chip are as the followings: computeraided design of microfluidic channel networks were developed using the AutoCAD software, and photo-plotted (printed) onto transparent masks. The schematic illustration of the PDMS stamp fabrication procedure is described in Fig Masters with a final feature height of 150 µm, that was verified by optical profilometry (Wyko model NT1100 Optical Profiling System, Veeco Instruments), were fabricated by spin-coating and subsequently pre-baking two consecutive layers of SU negative photoresist (Microchem,

70 Chapter 3. Experiment 55 Newton) at 1750 rpm on top of a seed layer of SU 8 25 that was spun at 2000 rpm. The transparent sections of a laser plotted transparency mask were selectively exposed to UV light, and subsequently developed and dried. PDMS was molded on top of the obtained master, then cured, peeled off, and cut to match the footprint of a 60mm 24mm 130 mm cover slip (No , Fisher Scientific). In order to bond the glass substrate to PDMS, the glass was treated with an oxygen plasma. The completed device was fluidically connected to the outside world using precut, deburred and passivated stainless steel pins as well as polymer tubing. All other specialty fluidic connections were obtained from Upchurch Scientific. Figure 3.13: Schematic illustration of the different steps of fabricating PDMS stamp. Modified from [93].

71 Chapter 3. Experiment 56 Current research methods for studying the blood vessel excised from mice that are not based on the microfluidic platform, have some disadvantages. These methods are expensive, time-consuming, and conducted manually which requires highly specialized training. Our collaborator (Guenther, Mechanical Engineering) has developed a polymerbased microfluidic device for loading and sealing the live blood vessel while also capable of using a pressurized flow system of MOPS buffer (physiologic salt solution) to keep the vascular tissue alive. Such microfluidic LOACs enable long-term study of vascular system without the problems and expense of previous methods that bring complex diagnostic tools into facilities for in vivo surgery. Fig illustrates the key features of this device. Resistance arteries were isolated by microdissection from 2nd or 3rd order mesenteries from wild type CD1 mice. An approximately 1 mm long resistance artery segment was introduced through a loading well, and was guided via pressure-driven flow (labeled A) to the inspection area. The future purpose of this design was to fix the artery spatially using flow-induced suction pressures (E) at eight points along the vessel. Subsequently, the artery segment was subjected to a microenvironment mimicking physiological conditions by simultaneously superfusing the outer vessel wall (B C) and perfusing the inner vessel lumen (A D). Figure 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].

72 Chapter 3. Experiment 57 The sample on which the experiment was carried out was the fresh dead blood vessel loaded to microfluidic chip containing MOPS solution. But it was not pressurized for the present preliminary investigation undertaken in this thesis study Fluorescent microspheres For calibrating the resolution of our purpose-built two-photon fluorescence imaging setup, images were recorded of 1 µm diameter fluorescent beads (Sigma-Aldrich,nile red fluorescent FluoSpheres beads) provided by Professor Axel Guenther group. The fluorescent particle was fluorophore manufactured from high-quality, ultraclean polystyrene. The carboxylate-modified microspheres were coated with a hydrophilic polymer containing multiple carboxylic acids for covalent attachment of ligands. The staining method places all of the fluorescent dye molecules inside each polystyrene microsphere instead of the bead s surface, which has a protective coating. This way the bead will be less affected by environmental effects that cause quenching or photobleaching of exposed fluorophore. The nile red fluorescent with excitation wavelength of 535 nm and emission of 575 nm have been used. The diluted microspheres (200:1) were placed on cover slides, and dried in an oven for 1/2 hr. The SEM and optical microscope images of the fluorescent bead is shown in Fig

73 Chapter 3. Experiment 58 Figure 3.15: Images of 1 µm diameter fluorescent beads (a) under SEM, and (b) under a bright field microscope with 100 objective lens.

74 Chapter 4 Results and discussion In this chapter, the experimental results are presented and discussed. First, we show the results of the two photon fluorescence imaging of 1 µm diameter fluorescent beads that assess the resolution of our imaging setup. Also, in this part we compare the results of two photon fluorescence imaging (TPI) with three different lenses and demonstrate the improvement of the cross-sectional resolution with increasing numerical aperture. In Section 4.2, two photon fluorescence imaging of a laser formed waveguide written inside bulk fused silica is shown. The imaging of the guiding regime in a single mode optical fiber follows in Section 4.3. The damage threshold data at the different interfaces of the microfluidic chip, including the cover slip surface, cover slip bottom interfaces with air, MOPS solution, and PDMS, and the bubble formation threshold in the MOPS solution presented in Section 4.4. It was necessary to identify the processing window in which the laser induced breakdown happens only in the blood vessel trapped in the microfluidic chip while avoiding damage or interaction with surfaces and materials in the chip. In addition, results of laser micromachining on the arterial wall in the form of a line cut and trepanning will be presented as a preliminary investigation. Finally, the combination of the two photon fluorescence imaging and the microma- 59

75 Chapter 4. Results and discussion 60 chining of the mouse artery will be given. 4.1 Two photon fluorescence imaging of microspheres The resolution of the multiphoton imaging setup was investigated by imaging 1 µm microspheres (section 3.4.2) deposited onto a microscope slide. The sample was translated in the x and y directions by using precise air-bearing motion stages. The fluorescent microspheres were doped with nile red fluorescent having an excitation wavelength of 535 nm and emission of 575 nm. As the laser central wavelength was 1045 nm, the laser was able to trigger two photon absorption in the microspheres. Two important factors should be taken into consideration for imaging. First, the power should be high enough to induce two photon absorption but not to cause photobleaching and photodamaging inside the sample. Second, the start position in the axial and lateral positions should be chosen carefully in the imaging process in order to be efficient in recording time. To find the proper laser power range for the imaging, the power varied from one tenth of a mw to a few mw while comparing the contrast and damage in the bead images. The best power was less than 1 mw to trigger the fluorescence signal in the fluorophore with low background noise. The power range would be different depending on the fluorescent material quality and its resistance to photobleaching. Moreover, the transmission of the lens should be considered as the power on the sample is an important factor. Another factor is the appropriate starting position. Each image takes a few minutes, hence, it will save a lot of time to know where to start scanning. By using the backlight and CCD camera (Sony XCD-X710), one can find the microsphere positions on the microscope slide. The surface of the cover slip was chosen as the axial start point. The surface of the glass can be recognized when the laser spot image is smallest and brightest on the CCD camera. In the cross section image (x-z), the start point is not critical as the z scanning range is much wider than the diameter of the sphere (for 1 µm scanning

76 Chapter 4. Results and discussion 61 step size it will be 256 times larger). In order to capture the top view image (x-y), the laser beam should be focused at the appropriate vertical position such that the spheres are in the depth of focus of the lens; in general it worked best if the Gaussian beam waist was changed to the center of sample (i.e. beads) being imaged. In taking the top view image, as the focus point is being scanned in one plane (z = z 0 ), the step size in both X and Y directions will be the same, while for the cross sectional image, the displacement of the focus point in the material is different from the lens displacement because of diffraction (section 3.3.1). In the following parts, two photon fluorescence imaging results will be given TPI by two dry lenses with different numerical apertures 40X-0.65 NA Images of fluorescent beads were recorded following the procedure of section and section 4.1. Briefly, a femtosecond laser at 1045 nm wavelength with 1 MHz rep. rate was focused on the sample using the NA lens with the transmission about 70%. The emitted fluorescence signals were recollected with the same objective lens. The power of the laser after the lens was set to 0.18 mw applying a ND filter. This amount of power was high enough to induce the two photon absorption in the sample but did not cause photobleaching, and hence the experiment was totally repeatable. Set the starting positions, the G-code was run to record the cross sectional image (x-z) of the fluorescent beads. From the result of this image, the microspheres center axial position could be determined. This axial position defined a suitable z position to take the top view image (XY). As an alternative way, one could take the top view image by simply focusing on the microscope slide, as the sphere size is just 1 µm which is within the depth of focus of the objective lens. Fig. 4.1 shows the result of the cross section image in which the axial intensity appears well-spread. Using ImageJ which is a public

77 Chapter 4. Results and discussion 62 domain, Java-based image processing program, the intensity profile of the beads in both directions was obtained. Average of axial Full Width at Half Maximum (FWHM) of the intensity profile (Maximum value/2) was found to be about 12 µm. To get a more accurate calculation of FWHM, the data was imported to MATLAB workplace in order to use the fitting tools. The Gaussian profile with the form of the a exp( ( x b c )2 ) represents the intensity data set and the calculation of diameter (FWHM) via d = 2 Ln(2) c gives similar values to similar to what was found with ImageJ. The value of d=11.73 µm was the average diameter found for 3 spheres. The axial intensity profile obtained by ImageJ and MATLAB fitting tools are compared in Fig Figure 4.1: Cross section TPI of the 1 µm fluorescent beads recorded with 40X-0.65 NA. The resolution of the laser scanning microscope depends on the optical components as well as the laser scanning step size. To capture the image, the step size was set manually. As a first step, the laser was focused on the sample surface and a step size of 0.5 µm was used while the G-code was run to record the x-y image as shown in Fig Then the step size was reduced to 0.25 µm to take another x-y image (Fig. 4.5). The improvement of the sphere image resolution as a result of the step size reduction is clear by the more spherical shape and the sharper image. The lateral intensity width (FWHM) calculation was repeated for 10 spheres to yield the value of 1.28 µm which is slightly longer than the 1 µm bead diameter. The Gaussian fit to the lateral intensity is illustrated in Fig. 4.6.

78 Chapter 4. Results and discussion 63 Figure 4.2: The axial intensity profile of one of a fluorescent bead represent in (a) ImageJ and in (b) MATLAB (red line).

79 Chapter 4. Results and discussion 64 Figure 4.3: Scanning laser beam through the microsphere beads mounted on the microscope slide. 100X-0.9 NA The same imaging procedure of section was repeated for the objective lens with 0.9 numerical aperture to compare the axial and the lateral resolution with what was obtained with the 0.65 NA lens. Knowing from the experiments with 0.65 NA that the power at the sample should be around 0.1 mw in order to trigger the fluorescence signal, and adjusting for the lower transmission of the 0.9 NA lens for infrared laser light which was about 50%, the optimal damage-free imaging power was found easily. Fig. 4.7 shows the x-y image when the step size was set to 0.25 µm. The average lateral intensity diameter (FWHM) over eight different fluorescent beads was 1.3 µm which was similar to the 1.28 µm value for the 0.65 NA case. This suggests that as long as the laser spot size (0.86 µm for 0.65 NA and 0.62 µm for 0.9 NA) be smaller than the 1 µm diameter of the beads, making the spot size smaller does not improve the resolution noticeably. In Fig. 4.8, the axial intensity profile (x-z) shows a 1.3x increase in the depth of focus resolution as the diameter (FWHM) was calculated to be 8.7 µm. This is a result of the larger NA that causes a shorter depth of focus for NA=0.9 in comparison with NA=0.65.

80 Chapter 4. Results and discussion 65 Figure 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.

81 Chapter 4. Results and discussion 66 Figure 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.

82 Chapter 4. Results and discussion 67 Figure 4.6: The lateral intensity profile of a fluorescence sphere (1 µm diameter) observed and represented by a Gaussian curve (red line) with MATLAB tools. In both cases of using 0.65 NA and 0.9 NA dry objective lenses, the axial length of the beads were much longer than the actual 1 µm diameter. This is due to the lens-like effect of the microspheres. As a result of the spherical shape of the bead, the laser beam will be focused as it enters the sample, so the actual focal position inside the bead is above the desired focus position. This spread the range of laser focusing axial point through the bead and lowers the axial resolution. The problem can be solved by using an oil-immersion objective lens which is presented in the next section TPI by oil-immersion lens Considering the smaller spot size (Eq. (3.3) and depth of focus (Eq. (3.2)) available for higher NA lenses, we found in Section that the lateral resolution of 1.3 µm obtained by two dry lenses (0.65 NA and 0.9 NA) was satisfactory, but the axial one which was 11.7 µm for 0.65 NA lens and 8.7 µm for 0.9 NA lens was not. In order to increase the axial resolution, an oil-immersion lens with 1.25 numerical aperture was tested here. The

83 Chapter 4. Results and discussion 68 Figure 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

84 Chapter 4. Results and discussion 69 Figure 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.

85 Chapter 4. Results and discussion 70 oil-immersion lens has the lowest depth of focus (1.93 µm) in comparison with the 0.6 and 0.9 NA lenses tested above. Also, the oil filled the volume around the microspheres dramatically decreases the refractive index difference between surrounding environment and spheres ( n < 0.12) in comparison with the dry lens where n = 0.6 difference exists between air and fluorescent beads refractive indicis. In this way, the lens-like effect of the beads will disappear. The fluorescent beads were soaked in oil with refractive index of , then sandwiched between a microscope slide and cover slip. The transmission of the oil immersion lens was lower than the two other lenses,namely, 30%; as a result, a lower ND filter was used to set an on-target power of 0.2 mw similar to the 0.65 and 0.9 NA exposure condition. A top view of the laser produced images of the microspheres are shown in Fig. 4.9 for a pixel size of 0.5 µm. The contrast of the image is lower in comparison with two dry lenses as a result of the lower collection efficiency of light from the oil immersion lens and the scattering of laser light from the oil surrounding the spheres. The cross sectional images are given in Fig for 0.5 and 0.25 µm pixel step sizes. The axial intensity width (FWHM) was calculated by ImageJ and averaged for over 30 spheres, to yield 2.1 µm width which improved around 5.6x and 4.1x comparing with the results for the 0.65 and 0.9 NA lenses respectively, which were 12 and 8.7 µm, respectively. The axial and lateral intensity profile FWHM for each lens calculated with ImageJ is outlined in Table 4.1. Lens specifications w 0 [µm] DOF [µm] axial length [µm] lateral length [µm] Nikon, 40X 0.65 NA Nikon, 100X 0.9 NA Nikon, 100X 1.25 NA Table 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.

86 Chapter 4. Results and discussion 71 Figure 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.

87 Chapter 4. Results and discussion 72 Figure 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. According to Table 4.1, the calculated beam spot size for all three lenses is less than 1 µm, this means that lateral laser images recorded with these three lenses will be approximately the same in aspect of resolution. Here, the collection efficiency and laser light scattering of the lens play the main roles in laser image contrast. Moreover, multiphoton absorption occurs in the length less than the depth of focus, however, in the case of spherical shape that has the lens-like effect, the recorded axial length is much higher than the depth of focus. Table 4.2 shows resolution calculation for three lenses using Rayleigh Resolution criteria definition where the lateral resolution is equal to 0.61λ/NA and the axial resolution is is equal to 2n oil λ/na 2. By decreasing the refractive index difference of fluorescent beads and surrounding environment the recorded axial value will decrease. Consequently, matching the refractive index of the oil and fluorescent beads will improve the results. Another factor is spherical aberration that affects resolution. By using the aspheric lens, one can get rid of this effect. The resolution for commercial confocal microscope is reported to be around 0.65 µm

88 Chapter 4. Results and discussion 73 for axial and 0.22 µm for lateral resolution (Zeiss LSM 510 META NLO). In two photon fluorescence microscopy the incident light wavelength is almost twice the one for confocal microscopy, so the resolution should be approximately 1.3 µm for axial and 0.44 µm for transverse resolution. With an oil immersion lens, it might be possible to get such a resolution in transverse direction with our purpose-built system. Lens specifications Calculated lateral resolution [µm] Calculated axial resolution [µm] Nikon, 40X 0.65 NA Nikon, 100X 0.9 NA Nikon, 100X 1.25 NA Table 4.2: Comparison of calculated lateral and axial resolution according to the Rayleigh Resolution criteria definition. 4.2 Two photon fluorescence imaging of the waveguides inside the fused silica As mentioned in Section 2.3.3, femtosecond laser imaging has a potential to be a tool for characterizing, analyzing, and visualizing optical waveguides inside a photonic circuit. Single mode waveguides were written inside fused silica by the laser exposure parameters listed in Table 4.3 which are similar to [2] and yield waveguides with 0.01 that guide at λ = 1550nm with 12.2 µm mode size. The objective is to demonstrate that such laser formed waveguides are detectable by two photon fluorescence imaging. To make it easier to find the waveguides for imaging, six ablation lines with 15 µm separation were created on the surface of the fused silica with 163 mw laser power scanned along the sample in the x direction. These lines were observable by eye. The laser polarization was in the x direction, so from literatures it is better to write the waveguide in the parallel direction (x) to get less insertion loss and smaller mode size. The objective lens applied for waveguide writing in the fused silica was a 40 aspheric

89 Chapter 4. Results and discussion 74 lens of 0.55 NA. Waveguides were written with 20 µm separation at 83 µm depth. For multiphoton imaging of the waveguides, pixel size was set to 1 µm to yield a 256 µm field of view along x. TPI was done with a 40X 0.65 NA lens with the power 2 mw at 522 nm wavelength and 1 MHz repetition rate. Both imaging and the writing were done with the same laser system and without any need to remove the sample. The only differences were the objective lens and the laser power. The imaging CCD camera with back lighting (Fig. 3.6) was used to find the position of the waveguides, consequently, the G-code program and SPC-830 software were run at the same time to scan the laser beam through the sample at these positions to capture the cross-sectional two photon fluorescence image of the waveguides. Rep. rate Scan speed Average power Wavelength Position Lens 1 MHz 0.75 mm/s 163 mw 522 nm 83 µm below the surface 0.55 NA Table 4.3: Laser exposure parameters for writing waveguides inside fused silica. The optical microscope and TP images of the six ablation lines on the fused silica surface and 28 waveguides 83 µm below the fused silica surface are shown in Fig. 4.11a and Fig b,c, respectively. The optical microscope image is an inverse of the real object but the TPI is not. The bright horizontal line line in Fig. 4.11c, which is the cross-section TPI of the sample, is the fused silica surface and the gap in horizontal line represent the ablation lines on the sample surface. Also, at 83 µm below the surface, the cross section of the waveguides are observable which are 18 µm long in the z direction and 3 µm in the x direction.

90 Chapter 4. Results and discussion 75 Figure 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.

91 Chapter 4. Results and discussion 76 The waveguide diameter in the x direction is higher than the laser beam diameter as a result of the heat accumulation effect which increase the heat affected zone. So, TP imaging of waveguides written inside fused silica is indeed possible. According to Fig which shows the cross section of a laser written waveguide, the laser beam will create two different modification regions with positive and negative refractive index variation regardless of the laser polarization. In TPI most likely only the positive refractive index change is observable. In order to record image of the negative refractive index change, one should increase the laser power well above the glass fluorescence threshold. This high power can induce damage in the sample. As the fluorescence signal can be both from the more dense material and the induced defects inside the laser interaction volume, there is other possibility that both positive and negative refractive index changes are observable in our TP image. The axial length of the recorded waveguide by TP image is 18 µm which is 1.5x larger than the axial length measured by phase contrast microscopic image (Fig. 4.12). Figure 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].

92 Chapter 4. Results and discussion Two photon fluorescence imaging of the optical fiber Another application of the two photon fluorescence imaging is for assessment and accurate probing of waveguides that are to be written in the cladding of the optical fiber. The Section 4.2 showed that laser formed waveguides inside fused silica glass are observable by TPI. TPI can be used to find the exact position of the core or the waveguide written in the cladding. For example, in writing a coupler it will be feasible to write a several waveguide at precise distances from existing one. In order to image the optical fiber by TPI, a 5-6 cm long segment of optical fiber was taped to a microscope slide (Fig. 4.13). Laser pulses with 1 MHz repetition rate at 522 nm with a power level around 2 mw were applied to the sample. Both the cross section and the top view images of the fiber were recorded with a 40X-0.65 NA and images are shown in Fig As the fiber surface is curved, it acts like a cylindrical lens. Consequently, in the cross section (xz) image, the fiber appeared to be longer than 125 µm diameter. Fig. 4.14b shows the top view image of the fiber optic, in which the 8 µm core is completely observable, but the fiber optic cladding interfaces are not visible. In order to trigger Fluorescence signal from cladding, the laser power should be increased. Figure 4.13: Single mode fiber optic taped on a microscope slide for TP laser image. In order to improve the axial resolution, a few drops of oil (Cargille 06350, n=1.4587) were put on the fiber optic sandwiched between the microscope slide and a glass cover slip. The oil has the same refractive index as the fused silica, therefore preventing the

93 Chapter 4. Results and discussion 78 Figure 4.14: TPI of a optical fiber with a 40X-0.65 NA dry lens (a) cross section (xz) and (b) top view (xy). lens-like distortion and spherical aberration of the fiber surface. The result of TPI is presented in Fig Oil has a fluorescence threshold lower than fused silica, and therefore, to keep the background fluorescence signal low, the laser power level should be chosen carefully. As shown in Fig. 4.15a, the fluorescence signal on the top of the cladding was emitted from the cover slip and the circular shape fluoresce signal of around 125 µm diameter is coming from the oil. Due to the existence of dirt and impurity inside the oil, the signal was not uniform. The core was expected to be observable in both top view and cross section image, but it did not appear in cross section image as shown in Fig. 4.15a. It maybe as a result of the contamination of the fiber or cover slip at the position which image was taken from and also the lower signal to noise ratio due to the lower laser power used here. The laser beam was scattered from the sample interfaces before reaching the core. More cross sectional data should be taken at different position of the fiber to confirm this assumption. Further, an image with a smaller field of view around the fiber should be recorded core position with higher laser power to improve image quality. In this case, as the coverslip and oil are out of the imaging area, the background noise will not be an issue.

94 Chapter 4. Results and discussion 79 Figure 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). 4.4 Breakdown threshold at various microfluidic chip interfaces As discussed in Section 2.2, femtosecond lasers can be applied to image and modify biological samples in order to create models of specific disease or to study the biological system functions. A part of thesis thesis was dedicated to combination of micromachining on an artery wall and two photon fluorescence imaging. The experiments were carried out on the 1-2 mm long segment of the artery loaded in the microfluidic chip as described in Section This work may be useful in the study of the angiogenesis, which will be explained in Section 4.5. However, for this purpose it will be necessary to keep the artery alive during the laser exposure. To keep the artery alive, it is necessary to apply pressurized flow system of MOPS buffer into the microfluidic channel. Because such a process required a special design for the microfluidic, this will be categorized as a future work. In order to get a rough estimation of the laser pulse energy required to damage the blood vessel tissue, a series of experiments were first done to measure the breakdown

95 Chapter 4. Results and discussion 80 threshold at different interfaces in the microfluidic chip. It was expected that the laser damage threshold of the live tissue would be below the damage threshold for these microfluidic chip components. It is essential to know the damage threshold of different microfluidic chip interfaces to evaluate the processing window of both the pulse energy and the number of pulses that induce damage only in the artery wall. On another note, bubble formation in the MOPS inside the channel was unwanted as for the future pressurized sample it will cause blood vessel death. So, measurement of the bubble formation threshold inside the MOPS solution was carried out. All of the experiments were done using the femtosecond fiber laser (IMRA µjewel) using a pulse duration of about 300 fs at 1045 nm with repetition rate of 1 MHz. Most of the present published work on the combination of surgery and two photon fluorescence imaging, uses two separate paths for either surgery or imaging whereas in this thesis, only one source and path was used. Also, for laser surgery two different regimes of the repetition rates and pulse energies have typically been applied to the sample: First, long series of pulses with repetition rates around 80 MHz and pulse energies well below the optical breakdown threshold; and second, low repetition rate around a few khz laser pulses with pulse energies slightly higher than the breakdown threshold. In the first regime, the heat accumulation effect is dominant as the time needed for heat to diffuse out of the focal point is around 1 µs in water. In the khz regime, each laser pulse independently induce damage in the specimen as the pulse energy is above the breakdown threshold [8]. In this thesis, 1 MHz repetition rate which is in the boundary of these two regimes was used. Further, at this repetition rate, our system produce up to 500 nj pulse energy which is much higher than the typicall fs oscillator output pulse energy of 10 nj. The transparency window of typical biological samples is in the infrared regime of µm (Fig. 2.6), while optical scattering coefficient scales with 1/λ. Therefore infrared laser light is more desirable for biological experiments and we selected laser wavelength

96 Chapter 4. Results and discussion 81 of 1045 nm to achieve deeper penetration in our samples Damage threshold: glass-air interface Top surface of the glass-air interface The first part of the experiment was dedicated to measuring the power needed to induce bulk changes in the top surface of the cover slip. The plan was to create parallel ablation lines on the glass top surface by variation of the laser power and the sample scan speed. In order to do so, the linearly polarized laser propagating in the z direction was focused on the sample surface through the X objective lens. The laser beam was scanned on the sample in the x direction in parallel lines. Each line corresponded to a different average power varied between 220 mw and 10 mw and different scan speed changing from 0.2 mm/s to 50 mm/s. The result of this experiment is shown in Fig For 80 mw power, and for scan speeds as low as 5 mm/s and below, ablation damage is observable; but for the 10 mm/s and higher scan speeds, only a bright line is induced which indicates a refractive index change inside bulk glass. For high scan speeds around 50 mm/s and 120 mw power, only a refractive index change is seen on the cover slip surface. According to what has been reviewed in Section 2.1.2, making an array of exposure also can be used to determine the breakdown threshold of the glass top surface. By varying the exposure parameters like power and number of pulses, one can use an optical microscope to recognize the damage induced in the transparent material. Fig shows the bright field microscope images of static exposures. The advantage of the static exposure is to get a lot of information in such a small area. For creating an array like this, it was necessary to use AOM to modulate the laser beam and control the on and off time of the laser in the time order of µs. Each point in the array is corresponding to specific number of pulses and average power. The exposure spots was separated by 20 µm in both x and y directions. Here, the average power was changed

97 Chapter 4. Results and discussion 82 Figure 4.16: Optical microscopic images of femtosecond laser line scan on the top surface of the cover slip interface with air. Each average power line scan is repeated for scan speeds from 0.2 to 50 mm/s. from 180 mw to 10 mw. The static exposure was sensitive to the lens focusing position as small as 1 µm vertically. In order to find the lowest threshold, the same laser exposure was repeated 18 times in different positions on the cover slip in varied focal positions of 1 µm steps, as illustrated in Fig In each one, the focus point defined by the CCD camera was added/subtracted by 1 or 2 µm. Among all data sets obtained from this experiment, four sets of data with the lowest number of pulses in each power that can induce observable damage on the sample were taken. Average and root mean square deviation (RMSD) of their number of pulses in each average power were calculated and the result is given in Fig Consequently, a laser pulse with a 150 nj pulse energy (150 mw power at 1 MHz repetition rate) is the threshold to induce damage on the glass surface. Also, the threshold number of laser pulses varied from 220 pulses at 110 nj pulse energy to 1 pulse at 180 nj pulse energy. Bottom surface of the glass-air The same static exposure of the previous Section was also carried out on the bottom surface of the cover slip against air and before it was bounded to PDMS (Fig. 4.20). The

98 Chapter 4. Results and discussion 83 Figure 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.

99 Chapter 4. Results and discussion 84 Figure 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. exposure was repeated 9 times and results showed less sensitivity to the focusing position in comparison with the top surface. From these 9 sets of data, 4 of them with the lowest threshold was taken for calculating the average and RMSD in order to draw the error bar for each average power and pulse energy. The results are shown in Fig laser pulses induced observable damage at 50 nj pulse energy while only a single pulse ablate the glass bottom at 120 nj pulse energy. The data in Fig which falling more rapidly than in Fig In Fig. 4.21, powers 80 mw regime only a single pulse is seem to have higher energy than the breakdown threshold. Consequently, even one pulse can create observable mark on the glass bottom surface. However, for power 70 mw one pulse energy is not high enough to induce damage and multiple pulses are needed to create damage i.e. 160 laser pulses at 50 nj pulse energy. In this domain a combination of heat accumulation effect and incubation of damage sites is playing a role. That lowers the pulse energy threshold. Fig shows the comparison between average number of pulses in each power

100 Chapter 4. Results and discussion 85 Figure 4.19: Minimum number of pulses corresponding to the specific pulse energy required 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.

101 Chapter 4. Results and discussion 86 Figure 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. needed to modify and damage the top and bottom surfaces of the cover slip. In a specific average power, the number of pulses needed to induce damage on the top surface of the cover slip is more than the ones needed to damage the bottom surface. One of the possible reasons is described in Fig The boundary conditions for the electric field at the first surface and second surface interfaces cause deenhancment in the electric field strength at the first surface and an enhancement at the second surface [95] due to a π- phase shift on reflection while a 0-phase shift for internal reflection at the second surface evaluated electric field. This will result in lower damage threshold at the glass exiting surface than the entering surface Damage threshold: glass-pdms interface Another interface in the microfluidic chip is between PDMS and cover slip. For variable number of pulses( 2000, 1000, 500, 200, 100, 75, 50, 25, 20, 10, 5, 2, and 1), and for the average power changing from 180 mw to 10 mw with steps of 10 mw, a static damage

102 Chapter 4. Results and discussion 87 Figure 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.

103 Chapter 4. Results and discussion 88 Figure 4.22: Comparison between laser damage threshold on the bottom and top surfaces of the cover slip for glass-to-air interfaces. Figure 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].

104 Chapter 4. Results and discussion 89 exposure was done for glass-pdms interfaces yielding the result shown in Fig For powers lower than 40 mw, no damage was observed under the optical microscope. Considering Fig. 4.25, pulse energy of 60 nj is the damage threshold for a single pulse and pulse energy of 40 nj is the damage threshold for multiple-pulse experiment. Consequently, the ablation threshold decreases with increasing number of pulses applied to the material as expected due to heat accumulation and incubation effects. Figure 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. This damage threshold tests were repeatable for most of the powers. The average and the error bar related to six sets of data are shown in Fig Damage threshold: glass-mops interface The same experiment as in Section was repeated for a cover slip-mops interface (Fig. 4.26). The MOPS solution is a kind of biological solution that is available in the microfluidic channel to keep the blood vessel fresh and to prevent the blood vessel from drying fast. Fig shows the damage threshold and error bars based on statistical calculation

105 Chapter 4. Results and discussion 90 Figure 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.

106 Chapter 4. Results and discussion 91 Figure 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. on five sets of data. For less than 60 mw power no interface damage was observed while 100 pulses at 60 nj pulse energy could induce damage. As a result, a single laser pulse with the energy of 90 nj can induce damage in the glass-mops interface Comparison between damage threshold results for different interfaces Fig compares damage threshold conditions for the cover slip top surface interface with the air and cover slip bottom surface interfaces the air, PDMS, and MOPS. The cover slip top surface damage threshold is the highest and the cover slip-pdms damage threshold is the lowest among all. As explained before, because of the electric field boundary conditions, the exiting surface has lower damage threshold than the entering surface of glass. Among the exiting surface interfaces interface with the air and cover slip bottom surface interfaces, the single

107 Chapter 4. Results and discussion 92 Figure 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.

108 Chapter 4. Results and discussion 93 pulse damage threshold varies from 60 nj for glass-pdms to 80 nj for glass bottom, to 90 nj for glass-mops, and to 180 nj for glass top surface. Further, all interfaces will undergo multiphoton damage at lower pulse energies with fairly steep pulse energy onsets. As PDMS is solid material, it will keep the laser heat at the laser spot position. Due to the lower heat diffusion time in PDMS in comparison with the MOPS and air, the damage process will be enhanced in the glass-pdms interface. On the other hand, MOPS solution has a cooling effect as it flows inside the channel. This, in turn, leads to cooling the glass down, and hence, an increase in the breakdown threshold in comparison with the glass-pdms interface. MOPS absorb laser heat more efficiently than air, consequently, MOPS-glass has lower damage threshold than bottom of glass-air. Table 4.4 shows the threshold pulse energies and their corresponding irradiance in single pulse and muliple-pulse laser exposure for different interfaces of the microfluidic chip. The irradiance was calculated considering 300 fs pulse duration, 1 MHz repetition rate and spot size equal to 0.86 µm (from Table 4.1). For each interface the damage threshold decreases as the number of pulses increase. This shows that for pulse energies lower than the single pulse shot threshold, a single pulse is not enough to induce damage in the material, and a larger number of pulses must exposed on the sample to induce damage via heat accumulation and incubation effects. Single pulse damage threshold Multiple-pulse damage threshold Pulse energy[nj] Irradiance[W/cm 2 ] Pulse energy[nj] Irradiance[W/cm 2 ] Glass-air top surface Glass-MOPS Glass-air bottom surface Glass-PDMS Table 4.4: Threshold pulse energy and irradiance for different microfluidic chip interfaces in single and a multiple-pulse laser exposure. Multipulses damage thresholds were found according to the end points in each line in

109 Chapter 4. Results and discussion 94 Figure 4.28: Comparison between damage threshold of different interfaces at microfluidic chip for various exposure condition.

110 Chapter 4. Results and discussion 95 Fig Bubble formation threshold inside MOPS solution To find the bubble formation threshold inside the MOPS solution, a primary experiment was done without controlling the number of pulses at 1 MHz repetition rate. The laser beam was focused inside the solution, and video was recorded using the backlighting and CCD camera. The rate of capturing the video was 25 frame/s and the CCD camera s field of view was covering the whole width of the channel (125 µm). At each exposure power, the laser beam was focused with X objective lens inside the MOPS solution while capturing video. The breakdown threshold in water obtained from this test was at 59 mw average power for 1045 nm 1 MHz 300 fs laser radiation. In the next experiment, for every desired average power, the MOPS solution was exposed with a specific number of pulses with a time interval of 2 second in which the laser was off. This time interval is longer than the bubble life time, so the bubble created by the previous number of pulses disappeared before new pulses arrived. Using this method, the bubbles were created only in the image plane of the CCD camera, and the bubbles larger than the CCD camera resolution ( 1µm) were observable. The result of the experiment can vary depending on the focusing position. For example, when the laser beam was focused near the PDMS surface, the threshold for bubble formation was lower as a result of the heating of the PDMS which can help the bubble formation or even generate gas from the PDMS. Consequently, the laser beam was focused about 30 µm below the cover slip in the microfluidic channel with 150 µm depth. According to [8], for the femtosecond laser there is a deterministic relation between irradiance and free electron density as shown in Fig Depending on the free electron density, various damage mechanisms can be involved in the laser interaction. According to the plasma density, mechanisms include breaking the chemical bond and formation of reactive oxygen, formation of nano-scale transient bubbles, and plasma mediated abla-

111 Chapter 4. Results and discussion 96 tion that creates a shockwave and long lasting bubbles [8, 96]. In our experiments, the irradiance was high enough to create long lasting bubbles with tens of millisecond life time. Pulse energy from 110 nj to 180 nj is corresponding to the irradiance between to W/cm 2. Figure 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]. Fig illustrates a plot of the number of pulses causing bubble formation in the MOPS versus the average laser power. The lines with different slopes are drawn as a guideline. The number of pulses are plotted logarithmically against a linear numbers for laser power. This data are based on averages from repeated experiments. At 120 nj pulse energy and less, the heat accumulation effect seems to be the main reason of bubble formation. In other words, a single pulse with energy in this regimes does not have enough energy to induce breakdown inside the MOPS solution. Also, in this regimes, by decreasing the pulse energy, more number of pulses is needed to trigger bubble formation and the rate of this increase is higher than the one for pulse energy above 120 nj.

112 Chapter 4. Results and discussion 97 The double slope in the bubble formation behavior suggests two different regimes in the nonlinear excitation. Figure 4.30: The minimum number of laser pulses inducing bubble formation inside the MOPS solution versus average power. The solid line is a guide. Some the bubbles formed after the laser exposure had a lifetime around 1 second. The lifetime of bubbles Depend on the power and the number of pulses. When there is a higher power than threshold power for a single shot, and the sample is exposed by multiple pulses the thermo mechanical effects of the laser exposure is dominant because of the both accumulation effect and high pulse energy. Fig combines data for the bubble formation threshold in MOPS and all previous damage thresholds. Eye-fitted curve has been drawn for the bubble formation threshold. This systematic damage threshold measurement under different exposure conditions was helpful the in laser exposure parameter choice. In this processing window, a laser micromachining can be done on the blood vessel without interacting with microfluidic

113 Chapter 4. Results and discussion 98 Figure 4.31: Comparison of all Breakdown thresholds at different exposure conditions. components.

114 Chapter 4. Results and discussion Micromachining on the mouse artery wall The growth and the formation of new blood vessels is called angiogenesis. Knowing the chemical and mechanical factors that can induce angiogenesis is important because this phenomenon happens in many diseases like cancer, arthritis, and psoriasis. In addition, angiogenesis is essential for some of the body functions like wound healing and maintaining adequate tissue oxygenation. In order to achieve better understanding of the angiogenesis, some studies have been done to detect the factors that influence and induce new formation of the blood vessel. One of the chemical factors that stimulate angiogenesis, for example, is vascular endothelial growth factor (VEGF) [97]. Other factors can also cause capillary formation, such as cutting a part of the live blood vessel with surgical scissors. Stimulation of the live blood vessel with femtosecond lasers is expected to induce capillary growth. To the best of our knowledge, no credible study has been carried out on this issue; hence, experiments confirming this belief are pointed out later on in the future work section of this thesis. Stimulating blood vessel capillary growth can be observed by the basic tissue interactions studies like: lysing of single cell, drilling through holes (trepanning), and cutting slits in the vascular walls. In order to become familiar with the micromachining on the biological sample, different experiments were done on the mouse artery wall. The artery sample had died a few hours before the experiment started. Although the MOPS solution in the channel helps keeping the blood vessel fresh, it was essential to work fast before artery dried and shrunk. As these samples were not fluorescently dyed, the only way to observe the laser damage on the blood vessel wall was using the bright field microscope and CCD camera. To make an observable mark on the artery wall, a high laser power was applied to the sample. The damage threshold of the biological tissue is well below the damage threshold of other transparent materials (Fig. 4.31). In our experiment, since the blood vessel was dead, it was necessary to make a visible mark on artery wall to be detected under the

115 Chapter 4. Results and discussion 100 Figure 4.32: Mouse artery wall loaded in the microfluidic chip. The laser beam ablated the blood vessel in the x direction. microscope before further laser damage study or TPI could be tested. For the entire artery machining experiments, the femtosecond fiber laser (IMRA µjewel) was operated at 1 MHz repetition rate with around 300 fs pulse duration at 1045 nm. The laser was focused with a 40X 0.65 NA objective lens. The AOM was applied to control the laser on and off time (Section 3.2). In the first set of the experiments, the laser beam was scanned in the x direction (Fig. 4.32) to cut across the artery. At each y position the same experiment was repeated 10 times. An x-direction laser cut was repeated for changing laser focus by step of 10 micron ( z = 10µm). This experiment was done for different powers and different scan speeds. The bright field microscope images are shown in Fig As the artery wall is not flat and smooth, some damage lines may be more observable in the different image planes. So, more than one image for each laser parameter is provided in Fig which shows the effect of power and scan speed on the tissue damage. For Higher laser power and slower scan speed, the laser cut is more visible and cut longer length because it was less sensitive to the initial focus position. Also, point ablation was created on the blood vessel wall by scanning the laser beam

116 Chapter 4. Results and discussion 101 Figure 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.

117 Chapter 4. Results and discussion 102 Figure 4.34: Scanning laser beam in the z direction (vertical to blood vessel surface) on the artery wall. in the z direction to pierce the wall. The damage induced by this experiment was hard to be found because of a small laser-tissue interaction point and roughness of the blood vessel wall. The damage induced by 100 and 150 mw average powers is shown in Fig Other sets of experiments were dedicated to trepanning on the blood vessel wall. The average powers of 80, 70, 60, 50, 40 mw were applied to the sample. The CCD camera and backlighting were applied to find the blood vessel surface. The beam spot was brought down slowly until the interaction between blood vessel and the laser was observed on the CCD camera. Then, from that position the G-code program for drilling the hole was run. The laser beam spot was moved down to 50 µm below the surface of the artery wall with the step size of 1 µm and the scan speed during the circular motion was 2 mm/s. Fig. 4.35a shows the image of a 30 µm hole drilled using 80 mw of power at 1 MHz repetition rate taken with the CCD camera and Fig. 4.35b shows the artery wall before and after the laser exposure by 70 mw average power femtosecond laser pulses. As mentioned in Section 2.1.2, in the time order of few picoseconds to tens of picoseconds the temperature in the focal volume increases via the the free electron energy thermalization.

118 Chapter 4. Results and discussion 103 As the time for the acoustic relaxation is higher than thermalization time interval, the temperature rise was confined in the laser focal volume to produce high pressure. This pressure will lead to the tensile force that can cause fracture of the material or in aqueous media leads to bubble formation. Usually the threshold for bubble formation is defined as a dissection threshold in cell surgery [8]. The bubble formation was observed for all of the laser powers tested except for 40 mw, which was not reproducible result. The bubble formation threshold of the tissue was lower than the one for the MOPS solution of 59 mw. The reason is that tissue is more absorbing (semi transparent; see Fig. 2.6) than the MOPS solution (transparent). Figure 4.35: 30 µm hole created by laser power of (a) 80 mw laser beam and (b) 70 mw. By using video editing software, some interesting frames were edited from the video captured with CCD camera during the trepanning and images are shown in Fig As the lens was moved down during the laser machining, the image plane was moved down as well. Bubble formation is observable in some of the frames like number 4 and 6.

119 Chapter 4. Results and discussion 104 Figure 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.

120 Chapter 4. Results and discussion Combination of trepanning and two photon fluorescence imaging In this section, experiments were carried out on only a limited number of dyed artery samples. This experiment utilized a combination of two photon fluorescent imaging and femtosecond laser ablation. The fresh unstained blood vessel can be also excited with the femtosecond laser and emit fluorescent light. The average power needed for unstained sample excitation is around 4-times higher than the average power needed for fluorescent dyed sample. Fig shows the two photon fluorescence image of the unstained artery taken by femtosecond laser pulses with around 7 mw average power. Note that the vessel here shrank and became smaller in the axial direction. Figure 4.37: Cross sectional two photon fluorescence image of the unstained artery wall. Scale is the same for both directions. Two types of fluorescent dye have been tested: Propidium iodide, and Fura/Fura-Red. The absorption wavelength of Propidium iodide is 532 nm (close enough to the one-half of the laser wavelength, 522 nm) and for Fura red is 472 nm which has a small overlap with one-half the incident spectral range. The results of fluorescence imaging with these

121 Chapter 4. Results and discussion 106 two different dyes are presented in Fig Figure 4.38: 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 (xy) view of the blood vessel which is dyed with Propidium Iodide. b) Cross section (x-z) image of the blood vessel with both Propidium Iodide (right image) and Fura-Red stained (left image). The only experiment was done in this area was the combination of trepanning on the blood vessel wall with the high power and two photon fluorescence imaging. First, with fluorescence imaging the surface of the blood vessel wall was found, then the G-code program for trepanning was ran at 1 MHz repetition rate femtosecond laser pulses with 160 mw average power. The G-code moved the lens down with the step size of 1 µm for 50 times. Fig shows the image of the blood vessel both before and after the laser exposure. Two photon fluorescence image of the blood vessel after trepanning is presented in Fig Due to the comparatively high laser power, the radius of the hole

122 Chapter 4. Results and discussion 107 ( 80µm) is larger than what was expected ( 50µm). Figure 4.39: The CCD camera images of the artery wall (a) before and (b) after laser exposure. The result of this experiment showed the potential of our experimental set up in performing TP imaging and laser micromachining by using one laser system. Laser trepanning with varied exposure conditions such as laser power combined with TP imaging will provide valuable data for working on the live blood vessel. Controlling a drilled hole on the artery wall is one step toward study of angiogenesis because using this method it is possible to create environmental change on the artery wall which may trigger capillary growth. Combination of the two photon fluorescence imaging with the laser ablation is helpful in finding the exact position of the sample. Also, with high resolution two photon imaging, it is possible to record the image of the sample before and after the laser ablation and compare the result to see the effect of the laser exposure on the sample.

123 Chapter 4. Results and discussion 108 Figure 4.40: Two photon fluorescence imaging of the artery wall after laser trepanning exposure. The diameter of the hole is around 80 µm.