investigation of structured fibres for nonlinear ... · declaration ii declaration i, navin prakash...

Investigation of structured fibres

for nonlinear endomicroscopy

imaging

A thesis submitted for the degree of

Doctor of Philosophy

by

Navin Prakash Ghimire

Supervisors

Prof. Min Gu

Dr.Hongchun Bao

Dr. Xiangping Li

Centre for Micro-Photonics

Faculty of Science, Engineering and Technology

Swinburne University of Technology

Melbourne, Australia

2015

Declaration

ii

Declaration

I, Navin Prakash Ghimire, declare that this thesis entitled:

Investigation of structured fibres for nonlinear endomicroscopy imaging

is my own work and has not been submitted previously, in whole or in part, in

respect of any other academic award.


Centre for Micro-Photonics

Faculty of Science, Engineering and Technology

Swinburne University of Technology

Melbourne, Australia

Dated this day, 22 April, 2015

Acknowledgments

iii

Acknowledgments

I am thankful to the Centre for Micro-Photonics at Swinburne University of

Technology for giving me an opportunity to pursue my Doctor of Philosophy. I

would like to thank my supervisors Prof. Min Gu , Dr. Hongchun Bao, Dr.

Xiangping Li for their effort in my research skills development. Prof. Min Gu’s

tireless effort and guidance have helped me conduct and convey scientific research

in a professional manner. I thank Dr. Hongchun Bao and Dr. Xiangping Li for their

suggestions and discussion in consolidating results.

I would like thank administrative staff and technicians for helping to conduct the

research smoothly. Thanks go to all my colleagues at the Centre for Micro-

Photonics for valuable discussion and suggestions during the period.

Finally, I would like to thank my parents, all my teachers, family and friends for

their encouragement and support.


22 April, 2015

Abstract

iv

Abstract

Broadband excitation and collection in a fibre-optic nonlinear endomicroscope is

realized by using a single hollow-core double clad photonics crystal fibre (DCPCF)

and a gradient index (GRIN) lens. Femtosecond pulses with central wavelengths in

the range of 750 - 850 nm can be directly delivered through the core of the fibre for

nonlinear excitation without pre-chirping. A gradient index lens with numerical

aperture of 0.8 designed to operate over the near-infrared wavelength range is used

for simultaneously focusing the laser beam from the fibre for nonlinear excitation

and collecting the fluorescent signal from samples. It is possible to optimally excite

different fluorescent markers with different excitation wavelengths without

externally adjusting a dispersion compensation for each wavelength using this

Abstract

v

system. This compact system is suitable to perform nonlinear imaging of multiple

fluorophors in the wavelength range of 750-850 nm.

In the contest of implementing super-resolution techniques such as stimulated

emission depletion (STED) in fibre-optic endoscopy imaging, beam propagation

for different input polarization states ( linear and cylindircal with/ without

superimposed the vortex phase ) through the core of the hollow-core photonics

crystal fibre (HC-PCF) fibre is characterized. Our experiment suggests that the

doughnut mode with the centre intensity null can be propagated through the core of

the hollow-core photonic crystal fibre. A doughnut mode and a fundamental mode

coupling can be achieved simultaneously in the core of the fibre by varying the

beam width and hence the effective numerical aperture (NA) through the single

coupling objective. This is a curcial step towards implementing the single

wavelength fibre-optic super-resolution endomicroscopic imaging.

Finally, realizing the need of a robust solid silica fibre to withstand the fast

scanning requirement of a miniaturized video rate imaging system and also with

the view of implementation of such fibre in a fibre based STED endomicroscope, a

double-clad solid silica fibre is newly designed for applications that need the high

efficiency operation of light at two colors. Ultra-short pulses at a central

wavelength of 800 nm are delivered by the core of the double-clad fibre which can

realize the transmission of the optical pulses with a net chromatic dispersion of

Abstract

vi

zero. This is achieved by integrating the double-clad fibre with a pair of long

period gratings, which allows optical pulses to propagate in a higher order mode

(LP02) in the middle of the fibre as well as in a fundamental mode (LP01) at the

beginning and the end of the fibre. The index profile of the double-clad fibre is

engineered so that the higher order mode has high anomalous dispersion that can be

used to compensate for normal dispersion of the fundamental mode. By controlling

the lengths of the fibre where pulses are in a fundamental and in a higher order

mode, the fibre with total zero dispersion can be realized. The new double-clad

fibre can collect 100% of visible light within the NA of 0.21 with a loss of the

optical pulses less than 1%. The design of this fibre is essential for applications

including fibre-optic nonlinear imaging for compactness, robustness, and low

optical power loss in dispersion compensation.

Table of contents

vii

Table of contents

Declaration ii

Acknowledgments iii

Abstract iv

Table of contents vii

List of figures and tables xi

Chapter 1: Introduction

1.1 Introduction 1

1.2 Thesis preview 7

Chapter 2: Background and literature review

2.1 Nonlinear microscopy 10

2.2 Optical property of tissue 13

Table of contents

viii

2.3 Fluorescence collection and imaging depth 16

2.4 Optical resolution 20

2.4.1 Super-resolution techniques 22

2.4.2 Single wavelength stimulated emission

and depletion microscopy imaging 27

2.5 Fibre-optic imaging system 28

2.5.1 Optical fibres used in endoscopy 30

2.5.2 Endoscope probe components 35

2.6 Ultra-fast pulse propagation 40

2.7 Dispersion in Optical fibre 41

2.7.1 Long period fibre gratings 45

2.7.2 Dispersion compensation using LPG 47

2.7.3 Long period fibre grating design 48

Chapter 3: Hollow-core photonics crystal fibre for

broadband nonlinear endoscopy

3.1 Introduction 52

3.2Advantages of HC-PCF in nonlinear endoscopy 53

3.3 Pulse width measurement through HC-PCF fibre 55

3.4 Broadband excitation and collection in

nonlinear endomicroscopy 58

3.5 Experimental results and discussion 63

3.6 Conclusion 64

Table of contents

ix

Chapter 4: Propagation of a doughnut beam through

hollow-core photonics crystal fibre

4.1Introduction 65

4.2 Experimental setups 68

4.3 Experimental results 69

4.3.1 Linear polarization states 68

4.3.2 Radial polarization states 70

4.3.3 Azimuthal polarization states 72

4.3.4 Circular polarization states 74

4.4 Conclusion 84

Chapter 5: Design of zero dispersive double-clad fibre

for operation of two color light

5.1 Introduction 85

5.2 Zero chromatic dispersion 88

5.3 Mode conversions 96

5.4 High efficient operation of two color light 100

5.5 Conclusion 102

Table of contents

x

Chapter 6: Conclusion

6.1 Thesis conclusion 103

6.2 Future work 106

Bibliography 107

Appendix

Dispersion relation -cladding modes 134

Author’s Publications 139

List of figures and table

xi

List of figures and tables

Fig1.1 Diagram illustrating nonlinear interaction of light with matter …………...11

Fig. 2.2: Refractive index and image profile of GRIN lens………………..…......37

Fig 3.1: Experimental setup for pulse width measurement with frequency resolved

optical gating (FROG); Laser: Ti: Sapphire laser (Spectra-Physics, Mai Tai HP,

~100 fs,80 MHz, 690-1040 nm), ND neutral density, OBJ 1 & OBJ2: 20x 0.25 NA;

XS1&XS2:3D fibre coupling stage; M1&M2: Ultrafast mirrors; FROG setup,

Swamp Optics, GRENOUILLE-008-50-USB)……………………………………55

Fig 3.2: (a) Spectral intensity and phase, (b) autocorrelation trace of (c) measured

and (d) retrieved pulse after the propagation through the 1.5 m length…………...56


xii

Fig. 3.3: Pulse width measured at the output of the 1.5 m HC-PCF for different

wavelengths at 100 mw………………………………………………………….57

Fig. 3.4: (a) Schematic diagram of the experimental set-up for a broadband

excitation and collection system for single fibre nonlinear endomicroscopy. PMT:

photo-multiplier tube. (b) Enlarged part of the probe consisting of a HC-PCF and a

GRIN lens. (c)-(e) Mode profiles of the HC-PCF for different wavelengths. (f)

Output power at different wavelengths for the input power of 20 mW through a 1.5

m HC-PCF……......................................................................................................58

Fig. 3.5: Log-Log plot of the two-photon-excited fluorescence intensity (If) versus

the excitation laser power (Ip ) of fluorescent beads for wavelengths of 760 nm,

810 nm and 850 nm. …………………………………………………………...…60

Fig. 3.6: Log-Log plot of the two-photon-excited fluorescence intensity versus the

excitation laser power for different lengths of the fibre. ……………………...…61

Fig. 3.7: (a)-(f) Two-photon fluorescence images of 1 μm fluorescent beads (scale

bar: 5 µm) ; (g) Lateral resolution (full width at half maximum) of 1 μm fluorescent

beads for the excitation laser power of 4.5 mW at the samples. (h)-(i) Two-photon

fluorescence images of 2 μm fluorescent beads and the Rhodamine B dye (scale


xiii

bar: 10 µm) at 800 nm (h) with emission filter Semrock – FF01-647/51 and ( i)

Schott – BG18. (j) Log-Log plot of two-photon-excited fluorescence intensity of

Rhodamine B dye versus excitation laser power at 800 nm………………..……..62

Fig. 4.1: (a) Schematic diagram of the experimental set-up for the characterization

of doughnut beams through a hollow-core photonics crystal fibre. ND: neutral

Density, OBJ: objectives, BM: beam manipulation, VA: variable aperture, HC-

PCF: zero dispersion hollow-core photonic crystal around the wavelength of ~807

nm, CCD: charge couple device, A: Analyser ……………………..……….…….68

Fig. 4.2: (a) Intensity ratio of dip and peak of the fibre mode for different

numerical aperture. (b)-(f). Mode profiles for the input linear polarization states

and at different angles of the analyser with respect to the vertical

axis……………………………………………………………….………………..70


numerical aperture. (b)-(f). Mode profiles for the input linear polarization states

superimposed with the the vortex phase and at different angles of the analyser with

respect to the vertical axis……………………………………………….…..……72


xiv


numerical aperture. (b)-(f). Mode profiles for the input radial polarization states

and at different angles of the analyser with respect to the vertical axis…………..73


numerical aperture. (b)-(f). Mode profiles for the input radial polarization states

superimposed with the vortex phase and at different angles of the analyser with

respect to the vertical axis ………………………………………………………..75


numerical aperture value. (b)-(f). Mode profiles for the input azimuthal

polarization states and at different angles of the analyser with respect to the vertical

axis ………………………………………………………………………..………76

Fig. 4.7: (a) Numerical aperture versus the normalized intensity ( Imin / Imax) for

an azimuthal polarization state overlapped with a the vortex phase. (b)-(f). Mode

profiles for the input azimuthal polarization state superimposed with the vortex

phase and at different angle of the analyser with respect to the vertical axis kept

before the CCD. ……………………………………………………….………….78


xv


numerical aperture. (b)-(f). Mode profiles for the input circular polarization states

and at different angles of the analyser with respect to the vertical axis

………………………………………………………………………..……..……..79


numerical aperture. (b)-(f). Mode profiles for the input circular polarization states

superimposed with the vortex phase and at different angles of the analyser with

respect to the vertical axis . (g) Doughnut mode profile at the fibre output, (h) cross

section intensity profile……………………………………………………………82

Fig. 5.1: (a) Schematic structure of a DCF intergrated with a pair of gratings for

achieving zero net chromatic dispersion. (b) Refractive index profile of the DCF.89

Fig.5.2: Mode evolution for mode LP01 and mode LP02 at wavelengths 725 nm

(blue), 775 nm (green) and 825 nm (red) for the w-DCF. The refractive index

profile is shown in background (black)……………………………………….…..90


xvi

Fig. 5.3: Waveguide dispersion Dw of a higher order mode in the DCF (a) for

different core diameters d1, (b) for different thicknesses of the index dip region d2,

and (c) for different cladding thicknesses d3. …………………………….…….92

Fig. 5.4: (a) Dispersion values for the higher order mode. (b) Dispersion values for

the fundamental mode…………………………………..………………..…...…93

Fig. 5.5: Total dispersion of the fibre under the condition (L1+L3) :L2 = 1:3.77…94

Fig. 5.6: Mode conversion of the first (a) and the second (b) long period grating

are the intensity of input in mode LP01, input in mode LP02,

output in mode LP01 and output in mode LP02……………………………………98

Fig.5.7: (a) Schematic diagram of the DCF for operating two color light. (b) The

fibre collection efficiency η of the visible light versus the NA of the beam coupled

by an objective………………………………..………………………………….101

Table 2.1: Comparison of DCF and DCPCF features for fibre endoscopes…….32

outoutininHFHF

IIII ,,,

Chapter 1

1

Chapter 1

Introduction

1.1 Introduction

Nonlinear imaging is one of the best optical imaging tools available today for

investigating cellular functions, the cell - cell interaction and the cell migration in

dense thick deep tissues. Deep tissue imaging is possible because of nonlinear

excitation with a long wavelength for achieving reduced scattering/absorption and

diffuse fluorescence collection from a spatially localized focus spot that limits

photo toxicity/bleaching in the out of focus region [1, 2]. Nonlinear endoscopy

with a small probe and a long thin optical fibre has more advantages in flexibility,

handiness, and compactness than bench-top imaging systems for biomedical and

Chapter 1

2

clinical imaging applications [3-5]. Miniaturized hand-held fibre-optic devices are

ideal for their flexibility and potential use in live suspects long term imaging

studies to be able to trace the development stages of different complication in

different body parts that are difficult to access with minimal invasion. Nonlinear

fibre-optic imaging has added the advantage of improved resolution and deeper

tissue penetration over other linear systems [5, 7]. Fibre-optic nonlinear imaging

devices are the only available tools for use in biomedical research for the three-

dimensional (3D) visualization of cellular layers, long term imaging studies and

minimally invasive clinical diagnosis and surgical procedures of difficult to access

areas of living and moving suspects.

Different fibre types have been employed in nonlinear microscopic imaging

systems over a decade to improve the image quality by efficient excitation and

collection of the fluorescence signal [8-11]. The double-clad fibre structure has

made it possible both to excite samples through the core and simultaneously collect

the fluorescence signal through the cladding of the fibre. The use of a fibre coupler

in the system makes it possible to separate the excitation and fluorescence signal

within the fibre [12-15]. Combination of the zero dispersive fibre, the double-clad

fibre, the fibre coupler and miniaturized probes with 3D imaging capabilities would

make it feasible now to have all fibre endoscopy. Till now single mode fibre

(SMF), solid silica double-clad fibre (DCF) and double-clad photonics crystal fibre

(DCPCF) have been used in nonlinear endoscopy that utilizes a single fibre both

Chapter 1

3

for excitation and fluorescence signal collection. These studies show room for

improvements in various aspects. Though there are fibre designs that impose zero

dispersion to the ultra-fast pulses in the near-infrared region there are no studies

available that use single fibre for both excitation and signal collection. In this work

we make an investigation of structured fibres for single fibre nonlinear

endomicroscopy.

Although one type of fibre is preferred over the other types of fibre, the major

hurdle with fibre-optic nonlinear imaging is still imposed by the group velocity

dispersion in fibres. When femtosecond excitation pulses propagate through a fibre,

they suffer chromatic dispersion from the fibre which broadens the pulses and

reduces the nonlinear excitation efficiency. Femtosecond pulses with a central

wavelength of 800 nm are widely used in nonlinear imaging because the peak of

two-photon excitation wavelengths of fluorescein and acriflavine solutions is near

800 nm. They are the only fluorophores that has been approved for use in human.

All solid silica fibres have a high normal chromatic dispersion at a wavelength of

800 nm. To solve this optical dispersion problem, a pre-chirp unit external to a

fibre is normally used in fibre-optic nonlinear imaging systems for the dispersion

compensation [12, 16-20]. A pre-chirp unit generally employs grating pairs,

prisms, chirp mirrors and acoustic-optics modulators (AOM) [21, 22]. By adjusting

the distance between mirrors/prisms/gratings, the angular orientation of these

components and the beam size, the different frequency components of a pulsed

Chapter 1

4

beam obtain different optical path lengths and thus compensate for the chromatic

dispersion in the fibre [23]. However, endoscopic systems with a pre-chirp unit are

not only cumbersome to operate but also limited for the dispersion compensation of

pulsed beams with different central wavelengths. In addition, the dispersion

compensation by grating pairs or prisms requires a laser beam to have multiple

reflections or transmissions through the units, which induce high loss to the laser

beam power [24, 25]. For this reason the design and operation of fibre-optic

nonlinear imaging systems is limited to a single near-infrared (NIR) wavelength

and broadband excitation and collection feature in a single fibre based endoscopy

device is not realized until recently.

Resolution at the diffraction limited value in a fibre-optic nonlinear imaging system

is another major challenge, which researchers around the globe are trying to tackle.

Resolution is directly proportional to the wavelength used and inversely

proportional to the numerical aperture (NA) of a lens used. Thus shorter

wavelengths and higher NA oil immersion objective lenses are used to resolve a

feature size approximately half of the wavelength. Biological imaging of a thick

tissue sample imposes restriction to use of a short wavelength and use of oil

immersion objective in a contact mode. Thus smallest feature size that could be

imaged under the in vivo condition of thin sample is limited until recently before

the availability of bench-top super-resolution imaging technology.

Chapter 1

5

To directly visualize the sub-cellular structure (smaller than the wavelength

dimension), fluorescence microscopy combines the specific anti-body/dye labeling

technique with light microscope system for achieving diffraction limited/unlimited

bio-imaging [26, 27]. Recently the optical nonlinearity of the fluorophores has been

used in combination to achieve diffraction unlimited imaging of fluorescent

labelled sub-cellular structures. Resolution beyond the diffraction barrier was

achieved by the combination of specific dyes/antibodies labelling of samples and

use of two lasers under the stringent alignment and synchronized condition [27-30].

Fibre-optic based endomicroscopic devices can offer the inherent alignment of two

beams for incorporating this advanced super-resolution imaging features like

stimulated emission depletion (STED) imaging. Detailed study of the propagation

of multiple wavelengths, multiple polarization states of optical-beams for

broadband excitation/emission is necessary to make these concepts feasible. Also,

using one wavelength for both two-photon excitation and quenching of the

fluorescence makes the optimization of optics of the STED system simple and

possible for 3D-imaging of thick scattering tissue samples.

In this thesis, we investigate the use of hollow-core photonics crystal fibre (HC-

PCF) in fibre-optic nonlinear endoscopy as well as theoretically design a DCF with

two values of the NA for the efficient transmission of two colors of light, near-

infrared optical pulses as well as the visible continuous wave (CW) bream. We

demonstrate the feasibility of broadband excitation and collection in a single fibre

Chapter 1

6

based nonlinear endomicorsocopic system using a piece of HC-PCF integrated with

a gradient index lens (GRIN). The DCF fibre designed to have two inbuilt long

period gratings (LPGs) enables the transmission of the near-infrared optical pulses

with zero net chromatic dispersion in its core as well as the high efficiency

collection of visible light coupled by an objective lens. This DCF is promising to

be used in a compact nonlinear fibre-optic imaging system.

Chapter 1

7

1.2 Thesis preview

Chapter 1 introduces the fibre-optic nonlinear imaging system, it’s applications and

explains the motivation behind the research works conducted along with methods

implemented. It also briefly describes the contents of each chapter.

Chapter 2 contains the theoretical background and review of the topics dealt in this

thesis. Topics like nonlinear microscopy and endoscopy, fibre-optic nonlinear

imaging, dispersion and its compensation in optical fibres, super-resolution

techniques and single wavelength super-resolution imaging are explained and

reviewed in detail in this chapter. The chapter starts with fundamentals of light

mater interaction and nonlinear microscopy. Optical properties of tissues like

absorption, scattering and their wavelength dependence and the propagation of

light through tissue are discussed in relation to imaging. Ways to maximize the

extent of excitation and collection from the depth for two-photon imaging are

explained in detail. Optical resolution, diffraction limits and super-resolution

techniques are explained in brief. Different types of optical fibres used in nonlinear

microscopy and different technical challenges like linear and nonlinear properties

of the optical fibre are described in relation to achieving the zero dispersion and

thus understanding the propagation of ultra-fast pulse at near-infrared wavelengths

Chapter 1

8

for the successful implementation and design of fibre-optic endomicroscopy

system. Various relations for the calculation of linear dispersion and nonlinear

parameters like the third harmonic distortion, dispersion length, and nonlinear

length are given. The imaging property of a typical GRIN lens is described in

detail as this component is used in our experiments.

Chapter 3 presents the advantages of HC-PCF in nonlinear microscopy,

experimental results on pulse width measurement through HC-PCF using

frequency resolved optical gating (FROG), linear threshold power level

measurement of different length of the fibre and demonstrate the feasibility of

broadband excitation and collection in a single fibre based nonlinear

endomicorsocopic system using a piece of HC-PCF integrated with a GRIN lens.

Chapter 4 explains the advantages and importance of fibre-optic based super-

resolution endoscopy and advantages of the single wavelength implementation of

such a system. This chapter contains the experimental study on the generation of

fundamental and doughnut beams at different values of numerical aperture (NA)

and the characterization of the propagation of fundamental and doughnut beams for

different input polarization states through HC-PCF which is a crucial step towards

implementation of such a system.

Chapter 1

9

Chapter 5 starts with the review of fibre types and addition of different features

over time for use in nonlinear endoscopy since beginning. The problem of

chromatic dispersion and drawbacks of using pre-chirp units in nonlinear

microscopic system are explained. Chromatic dispersion compensation using a

higher order mode fibre is discussed and the necessity of new fibre design that

meets the requirement for both nonlinear imaging at near-infrared wavelengths and

its possible implementation in fibre-optic based STED technology requiring the

operation of two color wavelengths is presented. The mechanism to achieve zero

chromatic dispersion within the fibre is explained. The complete design and the

nonlinear parameter calculation are given along with simulation results for the

optimization of fibre design. Features of the newly designed fibre and the potential

use in fibre-optic super-resolution imaging system are discussed.

Chapter 6 concludes the thesis with the future work.

Chapter 2

10

Chapter 2

Background and literature review

2.1 Nonlinear microscopy

The fundamental physics of light matter interaction is utilized in optical

microscopy for imaging. The induced dipole moment by the vectorial electric field

during the light-matter interaction is given by [28],

.......

3)3(

2)2()1( EEEP

, (2.1)

where )(i is the i

th order nonlinear susceptibility tensor. The first order

)1( is

responsible for absorption and reflection of light in materials. Linear microscopic

systems like wide field microscopy and confocal microscopy utilizes the first order

interaction of light with imaging samples. The second order )2( is responsible for

Chapter 2

11

nonlinear optical phenomena like the second harmonic generation while the two-

photon absorption, the third harmonic generation and coherent anti stroke Raman

scattering are due to the third order susceptibility )3( . Nonlinear optical

phenomena use multi-photon processes so the corresponding optical microscopy is

called multi-photon microscopy. The imaging techniques that use the nonlinear

interaction between light and matter are two-photon excited fluorescence (TPEF)

microscopy, second harmonic generation (SHG) microscopy, third harmonic

generation (THG) microscopy, coherent anti-stokes Raman scattering (CARS).

Apart from imaging application, other non-imaging applications of nonlinear

excitation that are used in biological research are multi-photon fluorescence

correlation spectroscopy (MP-FCS), multi-photon cross correlation spectroscopy

(MP-FCCS) [29-32].

Fig1.1 Diagram illustrating nonlinear interaction of light with matter [33]

Chapter 2

12

Development of fluorescent probes over the period of time plays a major role in

bringing the microscopy to the current state of development though both

endogenous and exogenous fluorochromes can be used in bio-imaging [26-27].

Current state-of-art fluorescence optical microscopy utilizes optical nonlinearity as

well as nonlinear response (saturation of fluorescence emission rate with intensity

of light) of fluorescent probes. Recently developed microscopes are capable of

providing information about subcellular structures, biophysical and biochemical

phenomena.

Among other available techniques, TPEF microscopy is the mostly used nonlinear

bio-imaging technique for a high quality thick tissue imaging in vitro/ in vivo. The

possibility of the two-photon absorption was shown theoretically in 1931 by Maria

Goppert. The first two-photon fluorescence microscopy is demonstrated in 1990 by

Denk et.al [34]. For two-photon absorption to occur, two-photons have to interact

with other particles within ~10-16

seconds. This gives two-photon absorption event

a quadratic relation to light’s intensity. This intensity square dependence gives the

three dimensional localization of the excitation as the excitation probability outside

the focal region falls off by the fourth power with distance along optical axis. Two-

photon imaging is superior to single photon imaging owing to the limitation of

absorption to small focal volume [35]. This confined absorption has the advantage

of inherent 3D optical sectioning effect, reduced out of focus photo-bleaching and

Chapter 2

13

improvement in signal to noise ratio in thick tissue imaging [36-39]. The use of

long wavelengths also gives deeper tissue penetration and less absorption being

closer to the optical transparency window of tissues [7, 40].

2.2 Optical property of tissue

It is very important to understand the detailed optical property of specific tissue

types under investigation for imaging and other light assisted treatments [40]. In

general, biological tissues are relatively more transparent over the optical window

of 1000 nm - 1400 nm. In animal tissue Haemoglobin, lipids, water is the main

absorber at the near-infrared (NIR) wavelengths range. Deoxyribonucleic acid

(DNA), haemoglobin, lipids, amino acids, structural proteins are the main absorber

at Ultraviolet (UV) and visible part of the spectrum. DNA and proteins are

generally auto-fluorescence under UV and visible excitation. There is about 30%

protein in soft tissues. Within 10 micron depth all the portion of excitation

wavelength at 6-7 micron will be absorbed in protein due to the excitation of amid

bond in this wavelength range. Scattering is another important phenomenon which

is mainly due to in-homogeneities of refractive index in tissue because of the

presence of heterogeneous mixture of molecular structures. The strength of

scattering of light in tissue depends on the size and mixture proportion of different

molecules. Scattering is termed elastic if the photon energy remains unchanged

Chapter 2

14

after deflection. Elastic scattering phenomena like Rayleigh and Mie are the major

scattering process in tissues. Rayleigh scattering is predominant phenomena for

particle size much smaller than wavelength. It is nearly isotropic scattering and

wavelength dependent.

Raleigh scattering from dipole scatter much small than the wavelength of light is

given by [41],

),cos1(8 2

24

24

0

R

NII (2.2)

where I is the Intensity of light scattered, R is the distance from the scatter, N is

the number of scatters, α is the polarizability, θ is scattering angle, is wavelength

of light .

Mie scattering is a major phenomenon for particle size comparable to wavelength

of light. It is not strongly dependent on wavelength and is forward directional

scattering. Cell components like mitochondria, lysosomes, vesicles, collagen fibrils

are responsible for Mie scattering in biological tissue while cell membrane are

responsible for Rayleigh scattering in a cell.

Different scattering parameters for Mie scattering in tissue are

Scattering cross section ( s) is given by

s= QsAs , (2.3)

where Qs is scattering efficiency , As is the area of scattering particles.

Chapter 2

15

Scattering coefficient ( s ) is given by

sss NQ , (2.4)

where sQ is the scattering efficiency , sN is the number density of scattering

particles.

Strength of scattering is a measure in terms of mean free path ( ls ) which is average

distance between scattering events. For cell size comparable to the wavelength

scattering is mostly directed in the forward direction. Scattering mean free path ( ls)

is given by inverse of the scattering coefficient. Anisotropy (g) is a measure of

forward direction retained photon after a single scattering event with θ as the

scattering angles of photon. Parameter Qs and g are dependent on the size of the

cell structures, the index of the cell and it’s medium.

Reduced scattering coefficient is given by,

ss g )1(' , (2.5)

This coefficient is the measure of cumulative effect of forward scattering events,

incorporating the scattering coefficient and anisotropy together. Other parameter of

importance for imaging in tissue is the physical path length and optical path length.

The ratio of the optical path length to the physical path length is 4 or greater than 4

according to the type of tissues. The propagation of the photon through tissues can

be simulated with the knowledge of these parameters using Monte Carlo, transport

theory or diffusion theory [42, 43]. Modeling of optical properties of soft tissues

Chapter 2

16

like in visceral organs are given in ref. [41], shows that refractive indices of all the

constituent cellular components should be considered to get accurate optical

properties of the tissue under consideration.

Various diagnostic and treatment process have been developed utilizing the change

in absorption, scattering by tissues for example the amount of scattering from tissue

may vary due to diseases conditions for example, reduction of the Nicotanimide di-

hydrogen phosphate (NADH) concentration in tissues is one of the indications of

the tumour tissue as there is consumption of a large amount of oxygen. The reduce

in auto fluorescence level from NADH in this situation has thus been utilized for

early detection of cancerous tissue. Laser assisted in situ keratomileusis (LASIK)

utilize the absorption of light source in the tissue as the primary mechanism for

treatment.

2.3 Fluorescence collection and imaging depth

The selection of wavelength for excitation of given fluorophore/s determines

penetration depth in nonlinear imaging. Fluorophores are chosen according to their

two-photon cross-section. It is shown that scattered infra-red photons are not

effective to excite fluorophores in TPEF systems [44]. This is an advantage in

Chapter 2

17

using long wavelengths as they are relatively less scattered, which can increase

excitation and may compensate the need of more photons for two-photon excitation

(TPE) at long wavelengths. For nonlinear optical microscopy like two-photon

imaging only non-scattered light from the focal volume is considered as signals.

The signal power follows Lambert-beer like exponential decline with imaging

depth z. Two-photon signal from depth declines as [33]

,)(/22

2ss lzl

z

PE eeF

(2.6)

where ls is the scattering length, F2PE is the fluorescence signal at depth z.

Available power at desired wavelengths and the pulse width are key parameters to

be considered in choosing laser system for nonlinear imaging. Almost all biological

tissues scatter light strongly but with localized nonlinear action in two-photon

absorption processes, multiply scattered signal photons can be assigned to their

origin providing image from several hundred microns deep of living tissue.

In TPE fluorescence imaging optimization of signal collection and detection

efficiency is the most important factor. For scattering tissues, ballistic fluorescence

photon becomes negligible beyond a few tens of microns deep from the surface so

the efficient fluorescence detection requires large field of view to collect scattered

light from the focal volume [32, 45]. Large field of view objective has low

magnification thus high numerical aperture (NA) with low magnification will have

particularly high collection efficiencies [44]. For TPF imaging from the deep

tissue, detection systems should be designed to collect the diffuse light.

Chapter 2

18

Fluorescence F generated in the focal plane decreases proportionally to the square

of the fraction of the ballistic excitation photons,

,)/exp(.

2

0

ex

slzPF (2.7)

Maximum imaging depth in TPEF microscopy is given by

),)(ln( max0max zP

TPlZ

ex

s

(2.8)

where is the collection efficiency of the system, Zmax is the maximum imaging

depth, ls is the scattering length for excitation in tissues, P0 is the average power

incident on the surface of the samples, α is the setup parameter that depends on

fluorophore properties as well as detector and shot noise, is the pulse duration , T

is the pulse period.

The probability of the number of absorption of the two-photon pair ( na) is given

by [34].

,2

)()(

222

2

2

excPp

avg

excpahc

NA

f

Pgn

(2.9)

where gp is a unit less factor that depends on the temporal laser pulse shape ( 1 for

rectangular, 0.66 for Gaussian , 0.59 for hyperbolic-secant-squared ),

)(2 exc is

the molecular cross-section, Pavg is the average power, NA is the numerical

aperture, h is the plank’s constant, c is the velocity of light in space, p is the

pulse duration, fp is the repetition rate. The extent of two-photon fluorescent

excitation (TPFE) can be manipulated by altering parameters like pulse width,

Chapter 2

19

pulse repetition rate, pulse shape and selection of suitable fluorophores and their

excitation with given NA of optics.

Two-photon excitation (TPE) cross-section for picoseconds and continuous wave

(CW) excitation beam are larger than TPE femto-second cross-section but local

heating effect of samples is the problem for considering pico-seconds and CW laser

to be used for TPE fluorescence imaging [39,46]. Fluorescence intensity If in two-

photon excitation of fluorophores depends on the pulse width and the pulse

repetition rate and hence the average power of the laser beam used. Linear region

in log – log relationship between the input power and the signal power is the

verification of occurrence of the two-photon excitation in fluorescence samples

during measurements.

Long wavelengths have the advantage of deeper tissue penetration and less

scattering in tissues but the diffraction limited spot size is comparatively bigger

than that for short wavelengths. Combination of super-resolution techniques and

deep tissue imaging using longer wavelengths has the possibility of deep tissue

super-resolution for 3D imaging from thick tissue biological samples with suitable

far-red fluorophore labelling.

Chapter 2

20

2.4 Optical resolution

Optical resolution of light microscope is its ability to distinguish the two point

objects in close proximity [47]. This is fundamentally limited by the wave nature of

light. A point object will be projected as blur spot of finite size by optical imaging

system called point spread function (PSF). Two points closer than full width half

maximum (FWHM) of the PSF will be difficult to resolve because of the

substantial overlap in their image. Resolution is given by ∆(x,y)= /(βn sin(α) in the

lateral direction and ∆z= β /n sin2(α) in the axial direction, where is the

wavelength , n is the refractive index of medium and α is the half angle of the

imaging system. With best optics for bio-imaging, this value comes between 200

nm - 300 nm laterally and 500 nm - 700 nm axially. It is not possible to directly

visualize the sub-cellular structure smaller than this dimension. Fluorescence

microscopy combines the molecule specific labelling technique with light

microscope for the direct visualization [48-53]. Nonlinear interaction of light with

matter is utilized in nonlinear fluorescence microscopy which gives the capability

of inherent optical sectioning for 3D imaging, deep tissue penetration with

improved signal to noise ratio and slightly improved resolution [54-56]. Further,

nonlinear response in time and space of the different fluorescence labels are

utilized in recently developed microscopy techniques enabling one to achieve

Chapter 2

21

resolution above by one order of magnitude over conventional microscopy

techniques. Factor of √β in spatial resolution and improvement in signal to noise

ratio by rejecting out of focus fluorescent background was achieved using the

pinhole detection scheme in confocal microscopy. Reduction in point spread

function (PSF) is achieved by nonlinear excitation in nonlinear microscopy. Near

symmetrical PSF was achieved by increasing effective system NA by using two

opposing objectives for excitation and detection in 4Pi and I5M microscopy.

Structured illumination microscopy uses higher harmonics in illumination scheme

to increase image resolution. These different approaches helped to extend the far-

field optical resolution but they are fundamentally limited by the wave nature of

light. Recently the nonlinear saturation of fluorophore is utilized to reduce the PSF

beyond the diffraction limit. It is demonstrated that with the saturation of emission

rates of fluorophores one can create emission distribution with sharp edges and

strongly confined non-emissive region that was not possible in linear systems. The

saturation of the upper transition state of fluorophores faster than its recovery is

achieved by high intensity laser beam matching the difference in energy of this

electronic transition. Exact shape of the emission distribution is manipulated by the

distribution of intensity. The most popular use of this technique is stimulated

emission and depletion (STED) microscopy which uses high intensity doughnut

shaped beam to limit fluorophore emission around the periphery of the diffraction

limited emission spot at the centre formed by another laser. Similarly the

nonlinearity in time is utilized for high resolution by particle localization methods.

Chapter 2

22

It is now possible to calculate the exact position of single emitters with their

centroid of emission pattern in samples with high labelling density. To get the

detailed structural information, the precise control of emission of individual

fluorescent labels within the diffraction-limited volume and separating their

emission in time are used to construct the high-resolution image. Using these

recently developed high resolution techniques it is now possible to uncover the

biological process at subcellular and molecular level but these techniques are only

available in bench-top systems in a research laboratory environment.

2.4.1 Super-resolution techniques

Far-field optical technique like fluorescence microscopy can now achieve higher

resolution which is comparable to near field microscopy like near field scanning

optical microscopy (NSOM) that can provide 20 nm-50 nm resolution. Three-

dimensional imaging with an optical resolution, as high as ~ 20 nm in the lateral

direction and 40-50 nm in axial dimension has been achieved by far field

microscopy [57, 58]. There are a number of super-resolution techniques currently

available in far field microscopy which can in general be categorized into methods

that sharpen the PSF of fluorophores using nonlinear responses of fluorophores

such as in stimulated emission depletion microscopy (STED), reversible saturable

optically linear fluorescence transitions (RESOLFTs), structured illumination

microscopy (SIM), saturated structured illumination microscopy (SSIM), nonlinear

Chapter 2

23

structured illumination (NSI), selective plane illumination microscopy (SPIM),

digital laser scanning microscopy (DLSM), or that localize the individual

fluorescent molecules such as in photo activated localization microscopy (PALM),

fluorescence PALM(FPALM), stochastic optical reconstruction microscopy

(STORM), fluorescence imaging with one nanometre accuracy (FIONA) [47-53,

59-74] . Some of these techniques are slightly different from each other and some

of them also use one or more techniques in combination to achieve high resolution

[75-77].

In brief, their working principles are described as follows. In SIM, samples are

illuminated with light pattern of different spatial frequencies (produced by using

gratings), that captures high frequency information from features sizes smaller than

wavelength. Fourier signal post-processing techniques are applied to shift this high

frequency information to lower frequency within diffraction limit. SIM can only

achieve a factor of two improvements in resolution below the diffraction limit. In

SSIM, which uses SIM with saturation of fluorescence emission, the resolution is

not fundamentally limited by the diffraction but by the level of fluorescence

saturation practically achievable. SSIM demonstrated 50 nm resolution in two

dimensions [78]. Nonlinear structured illumination (NSI) uses harmonic generation

from the saturated excitation state of the fluorescent in samples and mathematical

processing of these harmonics. NSI may not be suitable for live imaging as

described by the author [78].

Chapter 2

24

The imaging depth as well as resolution can also be optimized by simultaneous

temporal and spatial focussing of pulse. Good spatial and temporal focussing of

pulse used as described in ref. [79] can be achieved by using 4Pi technique that

uses two divided beams from a point source which are focussed at a point using

identical and opposing high NA objectives [80]. At common focal volume,

resulting illumination PSF is given by two coherent beam interference. The

fluorescence signal is collected and focused to a point like detector. 4Pi technique

achieve 4 times narrower central PSF maximum than that of conventional PSF in

confocal microscopy, but this narrowing is accompanied by side lobes which limit

resolution in such systems. Post processing techniques to delete these side lobes are

required during reconstruction of the 3D images. With simple de-convolution

algorithm as described in [81] axial resolution as 0.5 micron can be achieved using

two-photon excitation with 4Pi confocal microscopy setup. Spatial resolution also

depends on labelling intensity of fluorophores (ls = Vs/V) and the probe size and the

preservation of ultra-structures during the sample preparation. This technique

cannot be used in live cell imaging and is possible only for bench top SHG setup.

Fluorescence image is formed by spatial co-ordinates of fluorophores thus

determination if position of each fluorescent probe molecules in samples with high

precision form the basis of super-resolution by single molecule imaging. A single

fluorophore is visualized by its diffraction limited blur in its image but the position

of the fluorophore can be precisely known by the centroid calculation of intensity

Chapter 2

25

distribution within PSF. In a well distributed fluorescent samples with the

assumption of a photon emission from each individual fluorophore, the

localization precision is given by [57]

Δ localization ≈ Δ / √ σ, (2.9)

whrere, Δ localization is the localization precision, Δ is the PSF size and N is the

number of measurement of fluorophores. For multiple molecules within the

diffraction limited proximity the separation of signal in overlaped regions in

spectral domain or time domain is necessary. Use of photo switchable fluorescent

dyes or proteins and activating them in different points of time each fluorophores

can be precisely localized to reconstruct the images during processing. In practice

resolution of ~ 20 nm has been achieved experimentally.

STED microscopy overcomes the diffraction barrier for resolution by reducing the

effective size of focus by switching off the fluorescence ability of the fluorophores

in the outer part of excitation focus. This switching of fluorophore is achived by

forced transition between electronically excited state (ON) and the electronic

ground state of the fluorophore(OFF) by incidence of a photon that matches the

energy gap between lower excitation state and ground state other than the

excitation laser. The spectral range of detector is chosen in such a way that it could

not detect the wavelength of the emitted photon after the quenching process.

Chapter 2

26

FWHM of the region from which the fluorescence can be emitted in STED with

first approximation is given by [82].

,1sin2

)(

sIIn

Ir

(2.10)

where is the STED wavelength, n is the refractive index of medium, α is the

semi-aperture angles of the objective lens and I is the intensity maximum of

engulfing zero, and Is is the characteristic intensity required for reducing the

fluorescence ability by a factor of two for dye being used. In STED, raising the

STED laser power the saturated depletion region expands. I/Is is much larger than

unity but RESOLFTs using photo switched fluorescent FP595 and dark state is

achieved at much less laser intensity. Focus engineering is necessary to a produce

doughnut intensity pattern for the quenching beam. Sample preparation and choice

of suitable fluorophores and their properties are one of the critical design

parameters for imaging in STED [83]. STED can also be used for live cell imaging

using auto-fluorescent proteins in tissues. Auto fluorophores like Yellow

fluorescent protein, blue fluorescent protein can be used to get signal along with

other organic dyes like Tetra-Methyl-Rhodamine (TMR), and process like enzyme

mediated labelling with Oregon Green, or self-labelling tag, for live STED

imaging. Atto 532 , Atto647N are the most used dyes for STED microscopy. Auto-

fluorescence intensity and the change in the spectrum measurement can distinguish

normal and various stages of cancer cells at different internal organs [84]. For

example , the diagnostic algorithm based on the combination of the fluorescence

Chapter 2

27

peak intensity ratios of I-350/I-470 at 280 nm excitation and I-390/I-470 at 330 nm

excitation yielded a sensitivity of 100% [95% confidence interval (CI) 0.95-1.0]

and specificity of 100% (95% CI 0.90-1.0) [85]. Similarly, suitable fluorophores

are needed to implement TPF imaging at 1200 nm wavelength and make such

comparative study of fluorescence intensity ratio at this excitation wavelength for

the diagnostic purpose. An auto-fluorescent protein called Neptune, which is far-

red fluorescent protein with excitation peak at 600 nm or above have been, reported

recently [86]; that may be used for intra-vital imaging in mammals with two-

photon excitation at 1200 nm.

2.4.2 Single wavelength stimulated emission and depletion imaging

Requirement to use two laser beams and associated complex optical design,

inflexibility in selection of dye and high cost and difficulty in incorporating them

into other commercial microscopy system are some of the technical challenges

remaining for wide use of bench-top single wavelength STED microscopy. Also

the use of two different wavelengths in this system makes the imaging of the thick

scattering tissue samples impossible due to different amount of chromatic

aberration in optics and tissues. The use of one wavelength both for two-photon

excitation and quenching of the fluorescence makes the optics optimization task

simple as well as the use of STED system possible for 3D imaging of thick

scattering tissue samples. Recent works [87-89] shows the feasibility of using

Chapter 2

28

conventional dyes and single wavelength bench-top STED microscopy with

reduced system complexity for super-resolution biological imaging.

2.5 Fibre-optic imaging systems

Different design trials to build ultimate compact systems for bio-imaging in vivo

environment are going on. Each of the systems built so far has some advantages

and limitations not found in others. The core concept is to develop miniaturized

optical endomicroscope/endoscope with high resolution, large working distance

and wide field of view with video rate 3D image acquisition capability to collect as

much information as possible from internal organs of living suspects [90-91].

Combining two or three imaging modalities like two-photon absorption (TPA),

SHG, Optical coherence tomography (OCT), CARS and wide range of

fluorophores for simultaneous multi-colour excitation) have also been tried to get

more information, at the cost of increased complexity of the device. Imaging

endoscope contains different components for beam and signal delivery and

scanning mechanism [30, 92-94]. Fibre and micro-optics techniques are used to

make the system compact and mult-functional at the same time [31]. A well

designed fibre-optic imaging probe can make it possible to access internal parts in

minimum invasive manner and are suitable for clinical diagnostic and surgical

Chapter 2

29

procedures, which otherwise would have been a major procedure . These fibre-

optic imaging systems with suitable probes have inherent flexibility to be used with

living suspects for the long-term imaging studies in various fields to understand the

detailed progressive development stages of different diseases and other natural

phenomena. In this direction research has been focus in last two decades and fibre-

optical systems are now considerably reduced in their system size and increase in

their functionality. Fibres in imaging devices are basically used for the remote

delivery of light and the collection of signals. Optical endoscopy systems with

endoscopes probe of very small diameters (few millimetres) are most suited to

image hollow tissue cavities within living body. Different fibre-optic fluorescence

imaging modalities have been used to date like single-fibre confocal, dual-axis

confocal, fibre-bundle confocal, two-photon fluorescence , single-fibre two-

photon, fibre-bundle two-photon, multi-focal two-photon, dual-clad two-photon

systems. They have different sets of advantages and limitations as described in ref.

[95].

Technical problems like group velocity dispersion (GVD), self-phase modulation

(SPM), self-steepening arise when fibres are used for high energy ultrafast pulse

propagation required for non-linear imaging. GVD is due to different spectral

components of a pulse travel at slightly different speeds within the fibre core which

lead to the temporal broadening of the pulse. SPM is a result of the intensity

dependent refractive index of the fibre material. Self-steepening is a non-linear

Chapter 2

30

effect arising due to the intensity dependent group velocity. The primary cause of

temporal pulse broadening in fibre based system is GVD but with the increase in

the pulse intensity, non-linear broadening of pulse will also occur.

2.5.1 Optical fibres used in endoscopy

The step index optical fibre with an inner core of radius ‘a’ has 1 or 2% higher

index given as ‘n1’ then the outer cladding of lower index given as ‘n2’ and light

is transmitted in them by total internal reflection. Numerical aperture (NA) of the

fibre is given by

,)( 2/12

2

2

1 nnNA (2.11)

Optical fibres are characterised by their normalized wave number V given by

V= (βπ a σA)/ , (2.12)

Fibre supports single mode (LP01) if it satisfies inequality condition V< 2.405 or

few mode (LP01 and LP02) if V< 3.8317 (for step index profile) and multimode

else. Multi-mode fibre (V> 2.405) guides more than one spatial mode. Transmitted

modes in the fibre increase approximately quadratically (as a function of V) with

the increase in core the diameter ‘a’. Single mode fibre (SMF) deliver light

efficiently and also acts as a pinhole detector rejecting out of focus fluorescence

emission in two-photon fluorescence imaging systems [8].

Chapter 2

31

Photonic crystal fibres are special optical fibres and have periodic arrays of air

holes in silica glass medium. A photonic crystal fibre is superior to others due to its

capability to endless single mode operation, dispersion engineering and high non-

linearity [96-97]. Hollow-core Photonic band gap fibre (PBF) guides wavelengths

within the transmission band in the central air core. There is no problem of SPM in

this fibre as there is no material in the core. PBF does not rely on total internal

reflection for guiding light. PBF uses diffractive effects with wavelength scale air-

silica arrays. Periodic arrays of air-silica lattice create photonic band gap which

prevents light at certain wavelength range to propagate in the cladding. These

wavelengths are localized in core and transmitted. However higher mode dispersive

pulse broadening can be more severe than in conventional fibres so for fibre

lengths greater than few meters needs additional dispersion compensation for

higher order modes.

Another class of fibre with a silica core and air-silica micro-structures as cladding

called large-mode-area photonic crystal fibre (LMA-PCF). It has large core

diameter up to ~35 m. SPM effect is reduced and light guidance is ‘endlessly

single mode’ in such fibres. Higher mode dispersion is less severe in these fibres

than in hollow-air core fibres.

A double-clad fibre (DCF) has two claddings apart from the central core. They are

especially important in single fibre endoscopes as they can be excited through the

Chapter 2

32

core and collect signals using cladding with high NA. This feature has improved

the detection efficiency in nonlinear imaging modalities by two order of magnitude

compared to nonlinear imaging using single modes. Double-clad photonic crystal

fibre (DCPCF) has large-mode-area silica core, periodic array of air-silica micro-

structure with low NA as the inner cladding and solid silica the outer region as

outer cladding with high NA. This kind of double-clad photonic crystal fibre is of

particular interest as low NA cladding can be used for guiding ultra-short excitation

pulses with reduced SPM and the outer cladding with high NA can be used for the

efficient collection of fluorescence signal. Outer multi-mode silica-air interface

can have NA around ~0.6 for the visible light. Apart from dual excitation and

collection a double-clad photonic crystal fibre can be formed into a coupler which

can be used to separate the visible fluorescence from the infrared excitation. Fused

tapering method can be used to fabricate double -clad PCF couplers [13].

Table 2.1: Comparison of DCF and DCPCF features suitable for fibre endoscopes

SN Parameters

DCF DCPCF

1 Core size 3.6 micron 16 micron

2 Inner clad 125 micron 135 micron

3 Outer clad 250 micron 350 micron

4 NA core 0.19 0.04

5 NA clad 0.21 0.62

Chapter 2

33

6 Coupling efficiency

laser to fibre

91% 82%

7 Power at core 70% 38%

8 Degree of

polarisation

0.95 @18mW

0.78 @90mW

0.98@50mW

9 Endoscopic system

with lens at fibre tip

Focal length :740 micron ;

spot size : 9um @1310nm;

CSF(lf)=450micron

curvature 90micron

Field of view 2.5mm X 2mm

Probe size: 0.4mm

Lens NA 0.12;

Spot size : 6micron

@800nm

10 Cost wise less costly costly

11 Signal collection comparatively low efficiency

due to low NA

high efficiency due

to high NA

12 Mechanical property robust not so robust ( major

drawback for fibre

scanning imaging

systems)

13 Dispersion

has high dispersion (GVD :

45,127 fs^2/m @805 +/-

4.5nm ) over 60nm

bandwidth )

has low dispersion(

GVD of Photonic

band gap fibre (PBF)

-14622 fs^2/m)

Chapter 2

34

14 Coupler Coupler with polymer inner

clad (contact length 16cm)

have achieved 30% clad to

clad coupling of the visible

signal.

Coupler permits

separation of single

mode excitation and

multimode visible

signals at the

detection arm.

15 Consideration for

excitation at 1200nm

Long excitation wavelength

needs more power for

excitation and DCFs

effectively excite

fluorescence due to more

power carried in core.

16 After dispersion

compensation

In DCF de-polarisation effect

can be improved

A single cladding micro structure fibre with a small silica core yielding highly

nonlinear behaviours with interaction to light is being used positively to harness

SPM for broadening the spectrum of ultra-short pulses for the super-continuum

generation.

Chapter 2

35

Fibre bundles are also used in imaging. The fibre bundle consistng of up to

~100,000 individual step-index fibres are closely packed maintaining relative

arrangement through the length. They are able to transmit intensity image in

pixelated form. Laser scanning can be achieved at the distal end of samples by the

sequential illumination of the individual fibre in the bundle which can be taken as

advantage but reduction of lateral optical resolution (calculated as core-to-core

distance divided by optical magnification) is the main disadvantage of such system.

2.5.2 Endoscope probe components

SMF gives severe temporal and spectral broadening for ultra-fast laser pulses.

Their NA is low (~0.1) for the signal collection, which is far from wavelength

range mostly used for excitation in a nonlinear imaging system. Multimode fibres

could be used but they also lack the capability to focus to a diffraction limited spot

on samples for the efficient excitation without use of special techniques or other

micro-optics like use of gradient index (GRIN) lens. The use of fibre could make

the system compact but can’t achieve beam scanning without using other beam

scanning components like micro-electro mechanical system (MEMS). There are

ways to manage dispersion and reduce nonlinearity by the use of LMA-PCF or

hollow-core PCF for excitation alone but equal design consideration is needed for

back collected visible fluorescence signal. DCPCF is a good candidate for

Chapter 2

36

optimization of both excitation and collected signal. DCPCF has LMA core for

single mode excitation with reduced nonlinear effects, microstructure inner

cladding for multi-mode guidance of visible signals. The inner cladding is

separated by the web of silica bridges that are small in dimension than the

wavelength of light guided in the inner core. GRIN lens are used to focus the beam

from the fibre to a focus spot on samples.

GRIN lenses are sub-millimetre in size. GRIN lens uses a variable concentration of

dopant in glass to achieve characteristic parabolic refractive index profile given by,

)2/1()( 2

0 Arnrn

(2.13)

where n0 is the refractive index on the centre axis, r is the distance from the central

axis. GRIσ lens has no curvature to achieve focus instead it’s nearly parabolic

refractive index profile gives focussing effect if cut in proper size. The period of

the sinusoidal path (pitch of the lens is given by

2/LAp , (2.14)

where L is the lens length, √A is the gradient constant.

Adjustment of the distance between the fibre and GRIN lens and choice of pitch of

the GRIN lens in system are variables for achieving variable working distance and

focus spot size [94, 95]. Effective NA of the system and hence the optimization of

the two-photon fluorescence excitation in a single fibre based imaging system is

determined by the distance between three port coupler and the GRIN lens with

different pitch values [ 13].

Chapter 2

37

Fig. 2.2: Refractive index and image profile of GRIN lens [13].

For GRIN lens immersed in a refractive index medium nm( m= 1,2 ) , a laser beam

emerging from the fibre at a distance d1 to the GRIN lens surface can be focused to

an image distance d2 given by,

)cos()sin()sin()/()cos( 101

2

0211202 ALnnALAdnALAnnALdnnd , (2.15)

where n0 is refractive index on central axis , n1 & n2 are refractive index of the

medium, L is the length of the GRIN lens, d1 & d2 are object and image distance

respectively from GRIN lens surface, √A is the gradient constant .

on the opposite side of the GRIN lens. GRIN lens can have NA in range 0.46-1.

Recently the formation of lens on the fibre tip by the collapse of air holes has been

used to acquire images using nonlinear microscopy [98]. Use of SMF couplers,

which act as a low pass filter for signal in the visible wavelength region and

provide inherent confocal pinhole, has made it possible for all fibre nonlinear

imaging system. Photonic crystal fibre (PCF) couplers with double-clad fibre have

also been implemented in nonlinear optical endoscopy [17, 99]. Recently super-

continuum light for excitation has been generated within fibres using highly

Chapter 2

38

nonlinear property of fibres enabling simultaneous CARS and TPEF images [100].

DCPCF preserves linear polarisation in core but there is depolarization effect in the

inner cladding region due to micro-structures in large core area. So the

development of polarization states maintaining fibres is needed to implement SHG

imaging using double-clad photonic crystal fibres [101].

To implement second harmonic generation phenomena for imaging using fibre –

optics, the fibre and fibre couplers used should have polarisation preserving

characteristics. Measure of polarisation preserving feature is given by

degree of polarisation ( ) defined as

)(

)(

minmax

minmax

II

II

, (2.16)

If the value of is close to 1 polarisation at every 90o shows birefringence effect of

the fibre. Linear polarisation in fused single mode fibre couplers is preserved at

certain incident angles for both CW and pulsed illumination [65].

DCF exhibits a low nonlinearity and self-modulation effect. DCF is also found to

have degraded two-photon signal levels and polarisation of signal if used in SHG.

Instead DCPCF has higher threshold of the nonlinearity and higher degree of

polarization states maintenance. DCPCF is found superior over DCF for SHG

imaging and two-photon imaging [24]. Compact two-photon fluorescence

microscope based on single mode three port fibre coupler has been demonstrated.

Chapter 2

39

In such a system coupler behaves as low-pass filter that can deliver an ultra-short

pulsed laser beam with 38% coupling efficiency in the infrared range (770 to 870

nm) as well as collect fluorescence signal in the visible range at low (1%) coupling

efficiency with single mode profiles [8].

DCPCF coupler can be used to separate the visible fluorescence from infrared

excitation. Two lengths of DCPCF are twisted and heated by hydrogen flame with

flame size of approximately 10 mm and then draw gradually in the fused region.

The elongation length determines the splitting ration of the coupler formed and

mode coupling in the core and clad region. If the elongation length is longer there

will be multi-mode in both arms of the coupler so optimization is needed in the

fabrication process. Axial resolution and signal level is dependent on the gap

length between the fibre coupler and the back surface of the GRIN lens used as in

two-photon imaging system using the single mode fibre coupler and GRIN lens.

Axial resolution of 10 m was obtained in a two-photon imaging system with the

DCPCF and the GRIN lens [99].

Piezo-electric driven 2D-tilt mirrors near the samples are used to scan beam from

single fibre in portable microscopy. Micro-electro mechanical system (MEMS)

mirrors are another rapidly emerged technology that could be actuated electro-

statically or electro-thermally. The size of MEMS size ranges from 0.5-2 mm and

gives up to 30 degree angular rotation with a low voltage. Their fabrication is

Chapter 2

40

complex process involving sequential material etching and deposition. Probe is the

major component in endoscope imaging system. There are many designs of probes

available in literature with various functionalities [93, 102-103]. Figures in ref. [93,

102-103] show the mostly used and the most robust recent design. Resonant

scanning type probe has both fibre and imaging lens resonating together attached to

a cantilever to reduce possible aberrations. Because of the mechanical robustness

drawback of DCPCF, it cannot be used in such a resonant scanning probe system.

In Dual-axis confocal imaging system one SMF delivers excitation light and a

second SMF mounted at angles collects fluorescence from the overlapping region

of the two fibre apertures. To achieve variable focus, liquid lens that can alter the

focus on electro wetting are also tested.

2.6 Ultra-fast pulse propagation

The understanding of the propagation of ultra-fast pulse near-infrared wavelengths

in optical fibres is critical to successful implementation of fibre-optic imaging

systems. For ultra-short pulse propagation in optical fibres we have to consider

both materials and waveguide dispersion. The material dispersion effect comes

from the frequency dependent refractive index of the fibre material, silica in our

case. It can be approximated by Sellmeier equation.

,13

122

2

2

i i

i

B

An

(2.17)

where, Ai and Bi are Sellmeier coefficients and is the wavelength in micron.

Chapter 2

41

The waveguide dispersion is determined from the propagation constant ( ) of the

optical pulse, which is given by

( )=n(ω) ω / c , (2.18)

where ω is the frequency of the pulse and c is the velocity of light. The dispersion

coefficients for the pulse propagation are obtained by its Taylor’s series expansion

around carrier frequency ω0 given by � = + � − � + �! � − � + �! � − � + ⋯, (2.19)

where, j is the mode number. Coefficient 1 is related to the group velocity of

pulse in the fibre by = /�� where ng is the group index. The second order

coefficient β represents the second derivative of propagation constant with respect

to the wavelength i.e dispersion of group velocity and is responsible for the pulse

broadening. β has positive sign in normal dispersion regime and has negative sign

in anomalous dispersion regime. The waveguide dispersion in optical fibres is

given by � = − � � , (2.20)

The coefficient γ in the equation is called the third order dispersion (TOD). The

effect of TOD becomes important near zero dispersion of the wavelength in the

design. Pulse propagation in optical fibres is described in terms of two

characteristics length name as dispersion length (Ld) and non-linear length (Lnl). � = ∣ ∣ , �� = � , (2.21)

Chapter 2

42

Prominent effects during the pulse propagation are described in terms of the ration

of these lengths. When dispersion length is much shorter than the nonlinear length,

dispersion is prominent in the fibre. Similarly when the nonlinear length is much

shorter than the dispersion length the pulse propagation is nonlinear which doesn’t

affect temporal profile of the pulse but result in spectral changes in the pulse.

For pulses with a pulse width (To ) less than 1ps, which is the case for nonlinear

imaging using two-photon , high order nonlinear effects up to eleven have to be

consider in the Taylor series expansion and the simplified pulse propagation

equation with the normalized amplitude U is given as [104]

� + �� = ��6��′ � + −�� |�| � + � � |�| � − �� | |� , (2.22)

where, �′ = | | , � = 0 , �� = �0 is the absorption coefficient of the Silica.

The nonlinear parameters are calculated for our new fibre design in chapter 5,

which are required for the ultra-short pulse propagation in the fibre. Split step

Fourier method is used for the pulse propagation simulation. The split step method

calculates different values by splitting total length in small steps where dispersive

and nonlinear effects are assumed to act independently.

Chapter 2

43

2.7 Dispersion in optical fibres

When the short pulse passes through dispersive materials, the pulse width stretch

but the spectrum remains unchanged. In visible and NIR region, all materials have

positive dispersion (red frequency leads blue frequency). Dispersion is given in

unit “fsβ” (or for optical fibre in ps/km-nm). Positive dispersion can be offset by

adding negative dispersion using prism or grating pair in the path of the beam.

Dispersion in the optical path is the major problem in excitation and signal delivery

when using fibres as media. Pulse broadening resulting from the dispersion can be

estimated using

,))/(68.71( 2/122

inDinout (2.23)

where D is the total dispersion in femtoseconds squared. Typical bench-top multi-

photon microscopy system has 5000 fs2. In silica based optical fibre material

dispersion is around -70 ps/nm-km @ 800 nm and around -20 ps/nm-km @ 1200

nm. Increase in pulse width by dispersion decreases peak laser power on samples.

Total dispersion has different components. One of them is material dispersion,

which is due to the different values of refractive index imposed by the material for

different wavelength components of the pulse. The difference in time for the

different spectral components to reach the other end of the fibre i.e pulse

broadening is given by

Chapter 2

44

,0

0

2

0

22

0

d

nd

c

L

(2.24)

where n is the refractive index, c is the velocity of light, L is the length of the path ,

0 is the wavelength. Material dispersion (Dm) is given by,

,0

L

Dm (2.25)

Unit for material dispersion in optical fibre is given in ps/km.nm. Imperial relation

to calculate the material dispersion is given by Sellmeier’s equation (2.17).

Another dispersion mechanism that is present in single mode fibres defined by its

V number as given in equation (2.12) is the waveguide dispersion. Even in the

absence of material dispersion group velocity of the spectral component depends

on frequency (ω). The pulse broadening given by

,.)( 0

0

2

2

2

V

bVdV

dn

c

Lw

(2.26)

and the waveguide dispersion is given by

],)(

[2

2

21

dV

VbdV

nnD

c

w

(2.27)

The waveguide dispersion parameter inside the bracket of expression for Dw in

fibre can be obtained by the empirical expression given by following equations,

assuming weakly guiding condition.

Chapter 2

45

,)834.2(549.0080.0

)( 2

2

2

VdV

VbdV

(2.28)

both dispersion is given in combination by the expression

D = ,2

2

2

2

d

dc (2.29)

The fibre dispersion slope coefficient (S) is given by

S = D/∆ , (2.30)

The normal dispersion in the fibre and other optical components can be

compensated by the anomalous dispersion obtained by the grating pair for pre-

chirped pulse [105-107]. The problem in use of these gratings is the excitation

power loss reducing available power at the samples.

Fibre Bragg grating (FBG) can also be used for the dispersion compensation within

fibre but that will not serve purpose in imaging system due to the high reflection

[108]. Even the small amount of reflection could cause imaging artifices and not

permissible in biological imaging system. But for the case of transmission grating

like long period grating back reflection from grating is negligible.

2.7.1 Long period fibre gratings

Long period gratings (LPGs) are used as optical filters, gain flatteners, lens free

fibre to fibre coupler and for the dispersion compensation in optical communication

Chapter 2

46

systems apart from their use in different sensors [109-113]. They are used for light

coupling between two co-propagating modes in an optical fibre or wave guide.

LPG has pitch range in several micrometres to several millimetres range. Grating

created by acoustic wave along the fibre [114] or grating created by periodic stress

[115] have also been used for coupling light between two guided modes. In these

gratings, there exists a specific resonance wavelength at which the coupling

between the guided modes is the strongest.

Fibre gratings formed in the core of fibre by periodic modulation of refractive

index section that convert the fundamental mode present in the fibre into higher

order mode ( LP11 or LP02 ). LP01 to LP11 mode conversion is complicated because

of asymmetrical modes. However, LP01 to LP02 mode converter involve single

resonance being circularly symmetric mode for LP02. The spatial periodic (Λ)

modulation of refractive index in fibre core matches to inter-modal beat length

between the mode to be coupled given by the phase-matching condition.

,2

21

(2.31)

where 1 , 2 are the propagation constants of the two coupled modes.

At the wavelength where mode conversion occurs the optical power of the mode

LP01 is transferred to LP02 mode. With the decrease in power mode conversion

efficiency can be calculated in a transmission spectrum. The value of required

Chapter 2

47

refractive index difference ∆n for grating and fibre core radius rc for few modes in

the fibre are crucial design parameters.

LPG can be fabricated using femto-second laser beams in the core of fibres (which

are doped with Ge ) by inducing index modulation achieved by Ge and O2 bond

(defect) breakage [116-119].

2.7.2 In Fibre dispersion compensation using LPGs

The major problem of implementing all fibre based imaging system is the

dispersion compensation in fibres for femto-second laser pulses. In current

nonlinear imaging systems, prism and gratings pairs are used for the dispersion

compensation. It has been shown that normal dispersion in the fibre can be

compensated by anomalous dispersion achieved with long period grating in fibres

for higher order modes to make net zero dispersion in fibres for the wavelength

used. Nonlinear microscopic imaging system currently utilizes 800 nm as

excitation wavelength. Shifting the excitation wavelength to 1200 nm will increase

the penetration depth in tissues, but the system should be designed carefully within

available average power of laser sources to get two-photon fluorescence signals, so

dispersion issue will be more critical.

Chapter 2

48

There are methods for the dispersion compensation by converting the fundamental

mode to higher order modes and back to fundamental modes by using long period

grating at telecommunication wavelengths. There has been not much use of long

period gratings for the dispersion compensation at the wavelength of 800 nm for

endoscope imaging purpose. We propose the design of LPG for in-fibre dispersion

compensation at 800nm for two-photon endoscope imaging purpose.

2.7.3 Long Period fibre grating design

Resonance condition for fibre LPG is given by

,))()(( cladcoreres nn (2.32)

where ncore is the effective index of the fundamental core mode, nclad is the effective

index of the resonant cladding mode, Λ is the period of the core refractive index

modulation. The core and cladding effective indices can be expanded in Taylor

series about res. The impact of higher order derivatives in Taylor series for higher

order modes is less than 10% (error) to the value of core clad index difference at

the wavelength. For lower order modes (m<4) higher order values in the Taylor

expression contribute more errors and these values of dispersion should be taken

into account.The grating period that gives the coupling at the wavelength of p is

given by

Chapter 2

49

Λ= p / (neff-co –neff-cl ), (2.33)

where neff-co is determined from the dispersion relation to obtain the LP01 mode

normalized effective index b which is solved for the fibre from the expression

given by,

)(

)(

)1(

)1(1

0

1

0

1

bVK

bVKbV

bVJ

bVJbV

(2.34)

The effective indices of cladding modes can be determined by solving the

dispersion relation for three optical layers. Detailed dispersion relation derivation

for cladding modes is given in Appendix.

During LPGs fabrication, the refractive index difference achieved between UV

laser beams exposed portion and unexposed portion in the fibre core will determine

the grating characteristic like the coupling strength and the length of the grating.

The difference of effective index for the two co-propagating modes at chosen

resonant wavelength in the fibre will determine the period of the gratings required

for transfer of optical power between modes. Length of the grating is obtained from

the length required for the maximum power transfer condition. The amount of

anomalous dispersion achieved with this grating determines the distance between

the two LPG in series.

The length of gratings that gives the complete power transfer is

Lc= π/β , (2.35)

where the coupling coefficient ( ) between core mode LP01 and cladding modes

Chapter 2

50

with azimuthal value of 1 is given by [120] for step index fibres.

The imaging at the greater depth for available laser average power is limited by out

of focus fluorescence that is generated mainly near the surface of the sample [121-

124]. The ratio of focal (S) to out-of-focus fluorescence (B) is given by relation as

described in the manuscript of the fundamental imaging-depth limit in two-photon

microscopy is given by,

)/2exp()(2 2

2

s

s

lzznl

NA

D

S

(2.36)

where n is the tissue refractive index, ls is the scattering length.

For signal-to-noise ratio of unity (S/D= 1) at the excitation wavelength ‘ ’ of 1200

nm, tissue refractive index ‘n’ of 1.γγ and scattering length ‘ls’ of 390.65 m

(taking scattering coefficient of 2560 1m tissue at 1200 nm [125]) for NA of 0.5

and 0.21, imaging depths obtained are 1746.6 and 1290.2 respectively. These

values are two to three times greater than the 600 micron depth generally achieved

using two-photon microscopy at 800 nm wavelength.

This imaging depth can be improved further by simultaneous spatial focussing (by

focussing different wavelength component at same spot possible with aberration

free optics that uses two lenses with different materials) and temporal focussing of

pulse used as described in ref. [79]. Thus, if the dispersion is compensated then

using temporally focussed short pulses, penetration depth optimization is possible

in nonlinear microscopy [126-128].

Chapter 2

51

The STED system has been developed with common path for two beams from

different sources (excitation and depletion beam), which has made the STED

system as robust as previous; critical alignment of the two beams is not necessary.

The carefully designed phase plate (considering dispersion property of special glass

material for both beam at different wavelengths) produce required beam patterns

from the two beams which pass through it as described in ref. [129]. This setup was

able to observe dimension up to 60 nm.

Chapter 3

52

Chapter 3

Hollow-core photonics crystal fibre

for broadband nonlinear endoscopy

3.1 Introduction

A HC-PCF is unique for the use in nonlinear endoscopic imaging over solid-core

silica fibres because of its low chromatic dispersion, low nonlinearity, low loss and

a high damage threshold [130-132]. A HC-PCF is the most efficient fibre delivery

medium among fibre types, as the optical power in the HC-PCF propagates in the

air medium, which results in low transmission losses, low nonlinearity and low

Chapter 3

53

scattering [131,133-136]. In addition, this kind of fibre has no Fresnel loss in free

space coupling and gives a low background signal level [29,137]. The single mode

propagation at the 800 nm wavelength regime can be achieved in a HC-PCF

without sacrificing of the fibre core size. Further, a HC-PCF with zero group

velocity dispersion in the visible to near-infrared (NIR) wavelength range and

having large cladding numerical aperture (NA) with a large cladding diameter

allows us to have a high collection efficiency of the fluorescent signal in nonlinear

imaging.

In this Chapter, we demonstrate the feasibility of broadband excitation and

collection in a single fibre based nonlinear endomicorsocopy system using a piece

of hollow-core photonic crystal fibre (HC-PCF) integrated with a gradient index

(GRIN) lens.

3.2 Advantages of the HC-PCF in nonlinear endoscopy

Although the delivery of femtosecond pulses through a HC-PCF has been studied

for nonlinear excitation no endoscopy imaging has been achieved by collecting the

fluorescent signal through the cladding of the same piece of the fibre [102,138-

142]. Even for the investigation into a microscopy imaging system, the same piece

of the HC-PCF has not been used both for broadband nonlinear excitation and

Chapter 3

54

broadband fluorescent signal collection [143]. The collection of the fluorescent

signal in these systems is still conducted using a separate narrow band fibre other

than the HC-PCF [11,102 99, 138-142]. A single fibre endomicroscopy system

using a double-clad PCF (DC-PCF) has been developed both for excitation and

fluorescent signal collection but such a system requires a pre-chirp unit for

nonlinear excitation [16]. The ability to deliver the femtosecond light pulse in a

hollow-core over a broad range of wavelengths without the need for dispersion

compensation and the simultaneous collection of signal through the solid silica

cladding region gives the HC-PCF the advantage over other fibres used in single

fibre nonlinear endomicroscopy.

In this work, an endomicroscopy system formed by the integration of a single HC-

PCF and a GRIN lens is used. The system is tuneable over a broad wavelength

range of 750 – 850 nm without the need for the adjustment of the dispersion

compensation for each wavelength. The compactness of the system is improved by

avoiding the use of the pre-chirp units, which also reduces the optical power loss of

the system.

Chapter 3

55

3.3 Pulse width measurement through the HC-PCF fibre

Fig 3.1 Experimental setup for pulse width measurement with frequency resolved

optical gating (FROG); Laser: Ti: Sapphire laser (Spectra-Physics, Mai Tai HP,

~100 fs,80 MHz, 690-1040 nm), ND neutral density, OBJ 1 & OBJ2: 20x 0.25 NA;

XS1&XS2:3D fibre coupling stage; M1&M2: Ultrafast mirrors; FROG setup,

Swamp Optics, GRENOUILLE-008-50-USB).

Figure 3.1 shows the experimental setup for the pulse width measurement after the

propagation through 1.5 m HC-PCF. Figure 3.2 (a) shows the spectral intensity and

phase near zero wavelength of the fibre. The CCD image of the input pulse (Fig 3.2

(b)) within the frequency resolved optical gating (FROG) instrument and

Laser ND OBJ1

HC-PCF

M1 & M2

FROG

XS1

OBJ2

XS2

Chapter 3

56

algorithm retrieve pulse (Fig 3.2 (c)) give smooth the autocorrelation trace with the

FROG error of less than 1% as shown in Fig 3.2 (d).

Fig 3.2: (a) Spectral intensity and phase, (b) autocorrelation trace of (c) measured

and (d) retrieved pulse after the propagation through the 1.5 m length of HC-PCF.

(a) (b)

(c) (d)

Chapter 3

57

Fig. 3.3: Pule width measured at the output of the 1.5m HC-PCF for different

wavelengths at 100 mW

The output pulse width measured at the output end of the 1.5 m HC-PCF at the 100

mW input power by the FROG measurement is less than 100 fs for the wavelength

range from 750 nm to 850 nm as shown in Fig. 3.3. Therefore the compact

nonlinear endoscope as shown in Fig. 3.4(a) can be used for different wavelengths

without the requirement of the dispersion compensation. Thus, different bio-

markers can be efficiently used without the requirement for adjusting a pre-chirp

unit.

Chapter 3

58

3.4 Broadband excitation and collection in

nonlinear endomicroscopy

Fig. 3.4: (a) Schematic diagram of the experimental set-up for a broadband excitation and

collection system for single fibre nonlinear endomicroscopy. PMT: photo-multiplier tube.

(b) Enlarged part of the probe consisting of a HC-PCF and a GRIN lens. (c)-(e) Mode

profiles of the HC-PCF for different wavelengths. (f) Output power at different wavelengths

for the input power of 20 mW through a 1.5 m HC-PCF.

The schematic diagram of the experimental setup is shown in Fig. 3.4(a). It consists

of a Ti: Sapphire laser (Spectra-Physics, Mai Tai HP) which generates ultrafast

optical pulses with a pulse width of 100 fs, a repetition rate of 80 MHz and tuneable

Chapter 3

59

wavelengths around 690-1060 nm. Femtosecond pulses from the laser are coupled

through a 40x 0.85 NA objective to a HC-PCF (HC-800-01, NKT photonics, core

diameter 9.5 µm, core NA 0.2, cladding diameter 130 µm) with a zero dispersion

wavelength at 810 nm. For nonlinear imaging, a GRIN lens of numerical aperture 0.8

(GRINTECH, GT-MO-080-0415-810) is used in front of the output end of the fibre

for focussing the laser beam to the samples on a scanning stage and simultaneously

collecting the fluorescent signal back to fibre tip (Fig. 3.4(b)). The core and photonic

crystal regions do not support the visible signal as it is out of the band gap of the

fibre. Thus this part of the signal leaks out into the cladding region and propagates

with total internal reflection at the air and solid silica interface of the solid silica

cladding region after the outer acrylate coating of the fibre is removed. The

fluorescent signal is finally focused to a photo-multiplier tube (PMT). Band pass

filters (Schott - BG18 / Semrock - FF01-647/51) are used before the PMT to filter the

fluorescence signal from the reflected near-infrared light.

Figures 3.4(c)-(e) display the mode patterns in the HC-PCF at different wavelengths.

Since the wavelength 700 nm is out of the band gap of the photonic crystal in the HC-

PCF, the light could not be confined in the core, but distribute around the holey and

the cladding region of the fibre (Fig. 3.4(c)). On the other hand, the wavelength 800

nm is at the center of the photonic crystal band gap, the majority of light is confined

inside the core of the HC-PCF (Fig. 3.4(d)) and the remaining light leaks to the

cladding area. However, the wavelength 900 nm is close to the long wavelength edge

Chapter 3

60

of the photonic crystal band gap. Thus light confined in the core is decreased and

light leaked in the cladding increaseds (Fig. 3.4(e)) compared with the case in the

wavelength of 800 nm (Fig. 3.4(d)). Figure 3.4(f) shows the fibre output power

experimentally measured over the wavelength range for the given input power of 20

mW. Femtosecond pulses can be effectively delivered over a wavelength range of

750 - 850 nm through the core of the fibre. The transmission rates of the laser in the

fibre at the wavelength range is greater than 40% with highest 60% around

wavelength of 800 nm.

Fig. 3.5: Log-Log plot of the two-photon-excited fluorescence intensity (If) versus

the excitation laser power (Ip ) of fluorescent beads for wavelengths of 760 nm, 810

nm and 850 nm.

Dry fluorescent microspheres of diameters 1 μm and 2 μm (Fluoresbrite® Yellow

Green Microspheres, Polysciences Inc.) and Rhodamine B (Sigma Aldrich) were used

as test samples. The dependence of fluorescence intensity detected by the PMT (Fig.

(a)

Chapter 3

61

3.4(a)) on the excitation laser power at the samples for different wavelengths through

the HC-PCF fibre was measured. A plot of the fluorescence intensity against the

incident power on a logarithm scale for the central wavelengths of 760 nm, 810 nm

and 850 nm fits a straight line with a slope close to 2 (Fig. 3.5), indicating the

fluorescence signal detected by the PMT is due to the two-photon excitation for the

whole wavelength range.

Fig. 3.6: Log-Log plot of the two-photon-excited fluorescence intensity versus the

excitation laser power for different lengths of the fibre.

Figure 3.6 shows the plot of the two-photon-excited fluorescence intensity versus the

excitation laser power using dry fluorescent microspheres of diameter 1 μm as a

sample. For this experiment, the two-photon-excited fluorescence intensity for the

fibre with different lengths was measured by keeping the input coupling conditions

same. The threshold power where the fibre nonlinearity starts to play a role is

Chapter 3

62

different for different lengths of the fibre, as shown in Fig. 3.6. The threshold power

for fibre length of 200 cm, 140 cm and 80 cm are 75 mW, 92 mW, and 125 mW

respectively. These threshold power levels are higher than the corresponding power

levels that reported using other fibres [8,24] which is beneficial to the high efficiency

nonlinear endoscopy.

Fig. 3.7: (a)-(f) Two-photon fluorescence images of 1 μm fluorescent beads (scale bar: 5 µm) ; (g)

Lateral resolution (full width at half maximum) of 1 μm fluorescent beads for the excitation laser

power of 4.5 mW at the samples. (h)-(i) Two-photon fluorescence images of 2 μm fluorescent

beads and the Rhodamine B dye (scale bar: 10 µm) at 800 nm (h) with emission filter Semrock –

FF01-647/51 and ( i) Schott – BG18. (j) Log-Log plot of two-photon-excited fluorescence

intensity of Rhodamine B dye versus excitation laser power at 800 nm.

Chapter 3

63

3.5 Experimental results and discussion

Figures 3.7(a)-(f) show the two-photon-excited fluorescence images of the samples

containing the 1 µm diameter dry fluorescent microspheres over the wavelength

range, while Figs.3.7(h)-(i) show the two-photon-excited fluorescence images at 800

nm of the samples formed by mixing the 2 µm diameter fluorescent beads and

Rhodamine B fluorescent dye solution in water and dried on the glass slide. Emission

filters FF01-647/51, Schott BG18 are used to block the fluorescence from beads in

fig. 3.7(h) and from dye in fig. 3.7(i) respectively for imaging only the other.

Fig. 3.7(g) shows the lateral resolution (full-width at half maximum) for differrent

wavelengths. The slight degrading in resolution at short wavelengths is due to the

weak defocusing effect of the GRIN lens for the wavelength range. Fig. (j) Shows

the Log-Log plot of two-photon-excited fluorescence intensity of Rhodamine B

dye versus excitation laser power at the wavelength of 800 nm. It can be clearly

seen that Figs. 3.7(a)-(f) confirms the simultaneous boradband excitation and

collection (figs. (h)-(i) ), in the system shown Fig. 3.7(a), while Figs. 3.7(h)-(j)

demonstrate the multi-fluorophore two-photon imaging ability of the system.

Chapter 3

64

3.6 Conclusion

To summarize, broadband excitation and collection for a fibre-optic nonlinear

endo-microscopy system has been realized by using a single piece of the HC-PCF

integrated with a GRIN lens to operate in a NIR wavelength range from 750- 850

nm. The dispersion compensation adjustment for different wavelengths is not

required for nonlinear excitation over the range and thus multiple fluorescent

markers with different excitation peaks can be used simultaneously. The broadband

two-photon fluorescent signal can be collected through the cladding of the same

piece of the HC-PCF. The optical power level for system operation is reduced by

eliminating the requirement for lossy pre-chirp units for the chromatic dispersion

compensation. The HC-PCF has a higher nonlinear power threshold than other

types of fibres previously used in fibre-optic nonlinear endoscopy. The new endo-

microscopy imaging system is easy to operate, versatile, compact, and possible to

use low cost femtosecond pulsed fibre lasers.

Chapter 4

65

Chapter 4

Propagation of doughnut beams

through the hollow-core photonics

crystal fibre

4.1 Introduction

Different techniques have been available since a decade to achieve resolution

beyond long held diffraction limit barrier. Extensive reviews about different super-

resolution techniques and their applications can be found in a huge number of

literatures recently [144-150]. These systems are either hard to align, use complex

image restoration algorithm, or need specific laser/dye. All of them are limited to

bench top applications. The biological imaging community would tremendously

Chapter 4

66

benefit if these high resolution technologies could be implemented for routine in-

vivo imaging studies. With the use of optical fibre in biological imaging devices,

many bench top systems have been converted and are being used for clinical

diagnostic imaging and surgery; however the resolution of these devices are still

diffraction limited [3-5, 8-10,12,18-21,24-25,28, 99]. In resolution enhancement by

the spatial confinement of the excitation beam, for a scheme like in stimulated

emission depletion (STED) microscopy, phase plates are used on beam paths of

excitation and quenching laser beam to produce a doughnut shaped intensity profile

that overlaps with the central excitation spot. Resolution that could be achieved is

related to the dye property like the saturation intensity (Isat) and it’s rate(kqn)) and

the intensity of the quenching beam that could be achieved with a tight focus of a

high numerical aperture (NA) objective is given by the relation [144-150].

(4.1)

where λex is the excitation laser wavelength, n sinθ is the NA of objective used for

focussing two beams, the Isat value depends on the property of particular dye used

and Iqn is the intensity of the quenching laser beam achieved at the focus spot.

Previously fibre has been used in super-resolution imaging setup but its use is

limited mainly as the dispersion compensation element. Particularly the double-

clad fibre (DCF) structure is unique for use in nonlinear endoscopy as its different

layers can be used to excite dyes and collect the fluorescence signal simultaneously

,

sin2sat

qn

ex

I

In

res

Chapter 4

67

through the single piece of fibre [10, 16-18]. Recently, fibre-optical based super-

resolution imaging is demonstrated by using an azimuthal polarization state beam

instead of the use of the conventional doughnut beam for quenching the excitation

to achieved resolution beyond the diffraction limit in a fibre based endoscopic

imaging device [151]. With the first time demonstration of the concept, the system

can be further improved by using zero dispersive hollow-core photonics crystal

fibre (HC-PCF) and a single infra-red wavelength of for both excitation and

depletion. A single wavelength STED imaging system has also been successfully

demonstrated for bio-imaging at a wavelength of 770nm by using ATTO647N dye

[87-89, 152-154] in a bench-top system. Use of a single wavelength both for

excitation and quenching requires the implementation of pulse synchronization in

the time domain but simplifies the spatial overlap issues of two beams deep inside

the tissue owing the same optical response from components and tissues.

In this chapter we experimentally characterize the propagation of doughnut beams

through the core of a HC-PCF which is crucial for implementing compact single

wavelength single optical fibre based super-resolution imaging device for thick

biological samples. We use low NA coupling objectives to couple the light to HC-

PCF. The beam propagation condition for different input beam polarization states

and analysis after the fibre coupling objective are carried out.

Chapter 4

68

4.2 Experimental setup

Fig. 4.1: Schematic diagram of the experimental set-up for the

characterization of doughnut beams through a hollow-core photonics

crystal fibre. ND: neutral density, L1&L2: lens, HWP: half wave plate,

VPP: vortex phase plate, QWP: quarter wave plate, VA: variable

aperture, BM: beam manipulation, HC-PCF: zero dispersion hollow-

core photonic crystal around the wavelength of ~807 nm, CCD: charge

couple device, A: analyser

The schematic diagram of the experimental setup is shown in Fig. 4.1. It consists of

a Ti: Sapphire laser (Spectra-Physics, Mai Tai HP) which generates ultrafast

optical pulses with a pulse width of 100 fs, a repetition rate of 80 MHz and

tuneable wavelengths around 690-1060 nm. Femtosecond pulses from the laser are

coupled through a 10x 0.4 NA objective to a HC-PCF (HC-800-02, NKT

photonics, core diameter 7.5 µm, core NA 0.2, cladding diameter 130 µm) with a

Chapter 4

69

zero dispersion wavelength at 807 nm. Different input beam conditions are

achieved with different combinations of wave plates, a polarization state converter,

the vortex phase plates and a variable aperture in front of the fibre coupling

objective. The output of the fibre is imaged to the charged couple device (CCD) by

using the 40 x 0.65NA objectives. An analyser is inserted between the fibre output

and the CCD camera.

4.3 Experimental results

The HC-PCF can provide advantages over the fibre for implementing superresoltuion

imaging scheme. The input beam condition is not maintained at the output of the fibre

due to bifringence in the case of solid silica fibres, which is the major hurdle for

implementing STED scheme as circularly polarized beam superimposed with a phase

vortex are used to achieve a doughnut beam profile. The propagation medium at the

centre of the HC-PCF fibre is air and the propagation mechanism is fundamentally

different. HC-PCF can be designed to operate over a broadband at the designed centre

wavelength, which can provide an extra flexibility in choosing dyes arround 800 nm

for implementing STED.

The beam propagation condition through the HC-PCF for different input beam

polarization states with or without the superposition of the vortex phase and analysis

of the beam after the fibre coupling objective are given in the following sections.

Chapter 4

70

4.3.1 Linear polarization states without/with vortex phase

Fig. 4.2 (a) Intensity ratio of dip and peak of the fibre mode for

different NA. (b)-(f). Mode profiles for the linear input polarization

state and at different angles of the analyser with respect to the vertical

axis. Scale bar 5 µm.

0

0.25

0.5

0.75

1

0.1 0.2 0.3 0.4 0.5

Numerical Aperture (NA)

Imin

/Im

ax

(a)

(b) (c) (d) (e) (f)

Chapter 4

71

Figure 4.2 (a) shows the intensity ratio of dip to peak for linear polarization states.

Output beam profile from the fibre is imaged to CCD which changes from a

fundamental mode to a doughnut-like intensity profile with intensity deep in the

centre as NA of the input beam to the fibre input is increased. The ratio of

minimum intensity to maximum intensity at the centre achieved is around 0.5 for

this case. The ring shaped intensity profile breaks for different analyser angle with

respect to the vertical axis positioned between CCD and the fibre output as shown

in figure 4.2 (b)-(f).

Figure 4.3 (a) shows the intensity ratio of dip to peak for the case of linear

polarization states overlapped with a linear phase ramp obtained from the vortex

phase plate before coupling to the fibre input. Output beam profile from the fibre is

imaged to CCD which changes from a fundamental mode to doughnut-like

intensity profile with intensity deep in the centre as NA of the input beam to the

fibre input is increased. The intensity dip achieved is fast with change in NA for

this input polarization states. The ring shaped intensity profile breaks for different

analyser orientation to the vertical axis which is kept in front of CCD after the fibre

output as shown in figure 4.3 (b)-(f).

Chapter 4

72


different NA. (b)-(f). Mode profiles for the linear input polarization

states superimposed with the vortex phase and at different angles of

the analyser with respect to the vertical axis. Scale bar 5 µm.

0

0.25

0.5

0.75

1

0.1 0.2 0.3 0.4 0.5Numerical Aperture (NA)

Imin

/Im

ax

(b) (c) (d) (e) (f)

(a)

Chapter 4

73

4.3.2 Radial polarization states without/with vortex phase


different NA. (b)-(f). Mode profiles for the input radial polarization

states and at different angles of the analyser with respect to the vertical

axis . Scale bar 5 µm.

0

0.25

0.5

0.75

1

0.1 0.2 0.3 0.4 0.5


Imin

/Im

ax


Imin

/Im

ax

(b)

(a)

(c) (d) (e) (f)

Chapter 4

74

Figure 4.4(a) shows the intensity ratio of dip to peak for the case of radial input

polarization states. Output beam profile from the fibre is imaged to the CCD which

changes from a fundamental mode to doughnut-like intensity profile with intensity

dip in the centre as NA of the input beam to the fibre input is increased. In this

case, the fundamental mode intensity profile changes to the ring intensity profile

after redistribution of intensity for increase in NA. The doughnut-like intensity

profile after the analyser is achieved which breaks only for 45 º angle of analyser

to the vertical axis as shown in figure 4.4 (b)-(f).

Figure 4.5 (a) shows the intensity ratio of dip to peak for the case of radial

polarization states overlapped with the linear phase obtained from the phase plate

coupled to the fibre. Output beam profile imaged to the CCD camera change from a

fundamental mode to the doughnut-like intensity profile with intensity deep in the

centre as NA of the input beam to the fibre input is increased by opening the

variable aperture in front of the coupling objective. The intensity dip develops

slowly with change in NA for this polarization state. Orientation of side lobes are

changed for different analyser angle to the vertical axis between CCD and fibre

output suggesting radial polarization state at the output is change by the fibre as

shown in figure 4.5 (b)-(f).

Chapter 4

75


different NA. (b)-(f). Mode profiles for the input radial polarization

states superimposed with the vortex phase and at different angles of

the analyser with respect to the vertical axis . Scale bar 5 µm.

0

0.25

0.5

0.75

1

0.1 0.2 0.3 0.4 0.5


Imin

/Im

ax

(b) (c) (d) (e) (f)

(a)

Chapter 4

76

4.3.3 Azimuthal Polarization states without/with vortex phase


different numerical aperture. (b)-(f). Mode profiles for the input

azimuthal polarization states and at different angles of the analyser with

respect to the vertical axis. Scale bar 5 µm.

0

0.25

0.5

0.75

1

0.1 0.2 0.3 0.4 0.5


Imin

/Im

ax

(a)

(b) (c) (d) (e) (f)

Chapter 4

77

Figure 4.6 (a) shows the intensity ratio of dip to peak for the case of an azimuthal

input polarization state. Output beam profile from the fibre output imaged to the

CCD camera changes from a fundamental mode to a doughnut-like intensity profile

with intensity deep in the centre as NA of the input beam to the fibre input is

increased by opening the variable aperture in front of the coupling objective. The

fundamental mode intensity profile breaks into two side lobes for NA of 0.4. The

doughnut-like intensity profile is obtained after the analyser angle of 135 º to the

vertical axis on the CCD camera as shown in figure 4.6 (b)-(f).

Figure 4.7 (a) shows the numerical aperture verses the normalized intensity ( Imin /

Imax) for the case of an azimuthal polarization state overlapped with the vortex

phase obtained from the phase plate at the fibre input. Output beam profile from the

fibre imaged to the CCD change from a fundamental mode to azimuthal

polarisation states as NA is increased. Orientation of side lobe is changed for

different analyser angles to the vertical axis suggesting polarization states change

as propagated through the fibre as shown in 4.7 (b)-(f).

Chapter 4

78

Fig. 4.7 (a) numerical aperture verses the normalized intensity ( Imin /

Imax) for an azimuthal polarization state overlapped with the vortex

phase. (b)-(f). Mode profiles for the input azimuthal polarization state

superimposed with the vortex phase and at different orientation of the

analyser with respect to the vertical axis. Scale bar 5 µm.

0

0.25

0.5

0.75

1

0.1 0.2 0.3 0.4 0.5


Imin

/Im

ax

(b) (c) (d) (e) (f)

(a)

Chapter 4

79

4.4.4 Circular Polarization states without/with vortex phase



circular polarization states and at different angles of the analyser with

respect to the vertical axis . Scale bar 5 µm.

0

0.25

0.5

0.75

1

0.1 0.2 0.3 0.4 0.5


Imin

/Im

ax

(b) (c) (d) (e) (f)

(a)

Chapter 4

80

Figures 4.8- 4.9 show the beam propagation condition through HC-PCF for

different input polarization states without/with the overlapping of linear phase

ramp obtained from the vortex phase plate. Figure 4.8 (a) shows the intensity ratio

of dip to peak for circular polarization states. Output beam profile from the fibre is

imaged to the CCD camera which changes from a fundamental mode to doughnut-

like intensity profile with intensity deep in the centre as NA of the input beam to

the fibre input is increased by opening the variable aperture in front of the coupling

objective. The ratio of minimum intensity to maximum intensity at the centre is

achieved is around 0.5 in this case. The ring shaped intensity profile is maintained

for different analyser orientation to the vertical axis kept in front of CCD as shown

in figure 4.8 (b)-(f).

Figure 4.9(a) shows the fundamental and the doughnut mode at the core for different

effective coupling NA achieved by varying the beam aperture in front of the coupling

objective for the case of the circular polarization states superimposed with the vortex

phase achieved by the phase plate. The mode profile changes from the fundamental

mode to a doughnut mode with a centre intensity null as we open the variable

aperture allowing the higher order modes to couple into the core of the fibre. This

gives us the way to couple the differrent beam profile required for the STED to

couple to the core of the fibre using single objective by varrying the beam diameter of

the two beam path. Figures 4.9(b)-(f) shows the beam shape for the different analyser

angles to the vertical axis kept in front of the CCD. The doughnut beam with the

Chapter 4

81

central null intensity is maintained for any analyser orientation at the fibre output for

the input condition of circular polarization states with a superimposed vortex phase.

Figure 4.9 (g-h) displays the doughnut-like beam mode patterns and their cross-

section at the output face of the HC-PCF at 807 nm wavelength.

To implement a single wavelength optical fibre based STED system both the

fundamental excitation beam and the doughnut beam need to be propagate at the core

of the fibre to ensure spatial overlap of two beams at the sample plane as the

coupling of the doughnut beam coupled to the solid silica cladding could not maintain

its phase and polarization states in the silica medium.The spatial overlap of the two

beams at the sample plane at the output of the fibre is also not ensured if two beams

propagate through the physically displaced and different NA region of the fibre.

During experiments the different input beam conditions are achieved with different

combination of wave plates, polarization states converter, vortex phase plates and

variable aperture in front of the fibre coupling objective used for coupling the beam

to the core of HC-PCF fibre. Though the HC-PCF fibre are endlessly single mode for

and above the designed wavelength we use only short length of 1.5 m of the fibre to

achieve higher order propagation for different coupling conditions achivied by the

varriable aperture in front of the coupling objective.

Chapter 4

82



circular polarization states superimposed with the vortex phase and at

different angles of the analyser with respect to the vertical axis . (g)

Doughnut mode profile at the fibre output, (h) cross section intensity

profile. Scale bar 5 µm.

(a)

(b) (c) (d) (e) (f)

(g)

(h)

Chapter 4

83

For other cylindrical polarization states, the linear polarization states, azimuthal

polarization states with and without the superimposed vortex phase, the doughnut

beam shape is not maintained for different orientation of the analyser suggest that for

these input condition mode profile is not pure doughnut but combination of higher

order modes as shown in figures 4.2-4.7.

Comparing figures from 4.2 to 4.9, minimum intensity dip in the centre of

doughnut-like intensity profile achieved after propagation through the fibre is

deeper in case of addition of phase ramp using the vortex phase plate than without

for the circular polarization states. The doughnut intensity profile is maintained for

different orientations of analyser angles with respect to the axis for the both case.

Chapter 4

84

4.4 Conclusion

We characterized the beam propagation for different input polarization conditions (

linear, cylindrical and circular with/ without a superimposed vortex phase ) through

the core of the fibre. Our experimental results suggest that the doughnut mode with

centre intensity null can be propagated through the core of the HC-PCF for circular

polarization states superimposed with the vortex phase. Doughnut mode and

fundamental mode coupling can be achieved simultaneously to the core of the fibre

by varying the beam width through the low NA single coupling objective, which is

the crucial step towards implementing the single wavelength fibre-optic super-

resolution imaging endoscope.

Chapter 5

85

Chapter 5

Design of zero dispersive double-clad

fibre for two color light

5.1 Introduction

A single mode fibre (SMF) was first tried in nonlinear endoscopy for the delivery

of ultra-short pulses to specimens and the collection of nonlinear signal [24].

However, an SMF has limitation in fluorescence signal collection due to its small

core diameter. A double-clad photonic crystal fibre (DCPCF) was sought to

increase the collection of nonlinear signal due to its large inner cladding size [9, 24,

99]. A DCPCF has been reported to have a signal collection efficiency improved

Chapter 5

86

by two orders of magnitude compared to an SMF [9]. Despite the improvement in

signal collection, a DCPCF is mechanically less stable due to its air hole based

structure. System robustness is one of the critical requirements for nonlinear

endoscopy systems where the fibre is mechanically scanned and bent frequently

during imaging. To overcome those drawbacks, a double-clad solid silica fibre was

used in nonlinear endoscopy [10,25]. It has been reported that a double-clad fibre

(DCF) is not only mechanically robust, but also has high nonlinear signal collection

efficiency due to its larger inner cladding diameter [12, 16-18,155]. In addition, a

DCF has higher light confinement than a DCPCF. As a fibre coupler is integrated

in a nonlinear endoscopy system, it can further improve the system compactness

and the signal collection efficiency [23]. The signal collection efficiency of a DCF

coupler is one order magnitude higher than that of a DCPCF coupler.

In spite of these advances in nonlinear endoscopy, fibre based nonlinear endoscopy

systems face a common problem. When femtosecond excitation pulses propagate

through a fibre, they suffer chromatic dispersion from the fibre which broadens the

pulses and reduces the nonlinear signal excitation efficiency. Femtosecond pulses

with a central wavelength of 800 nm are widely used in nonlinear imaging because

the peak of two-photon excitation wavelengths of fluorescein and acriflavine

solutions is near 800 nm. They are the only fluorophores that has been approved for

use in human. All solid silica fibres have a high normal chromatic dispersion at a

wavelength of 800 nm. Currently, nonlinear endoscopy uses grating pairs or prisms

for dispersion compensation of optical pulses with a central wavelength of 800 nm

http://en.wikipedia.org/wiki/Fluorophore

Chapter 5

87

[3-17, 24-25, 99,156-158], but these units are bulky and are not stable in alignment.

In addition, dispersion compensation using grating pairs or prisms requires a laser

beam to have multiple reflections or transmissions through the units, which induce

high loss to the laser beam power [18, 25].

A higher-order-mode (HOM) fibre has been reported for chromatic dispersion

compensation of optical pulses within a fibre [159-164]. However the HOM fibre

has a single core/cladding structure and only uses its core for transmitting optical

pulses with a single numerical aperture (NA) determined by the core/cladding

index difference at a given wavelength [165]. The core of the HOM fibre supports

multiple modes and the HOM fibre needs to splice two single mode fibres at each

end of the HOM fibre to balance the group delay dispersion, which induces

additional loss. Therefore, it is limited in many applications including nonlinear

endoscopy which requires the operation of two wavelengths of light with different

NA.

In this chapter, we explain our a new DCF design with two values of the NA for

the efficient transmission of two colors of light, near-infrared optical pulses as well

as visible continuous wave (CW) light. The DCF with two inbuilt long period

gratings (LPGs) enables the transmission of near-infrared optical pulses with zero

net chromatic dispersion in its core as well as the high efficiency collection of

visible light coupled by an objective lens. This DCF is promising to be used in a

compact nonlinear fibre-optic imaging system.

Chapter 5

88

5.2 Zero chromatic dispersion

A solid silica DCF is mechanically robust. However, a fundamental mode at a

wavelength of 800 nm is far away from the zero dispersion wavelength (1270 nm)

of a silica fibre and it has high group delay dispersion of ~ -110 ps/nm·km [166].

As femtosecond pulses with a central wavelength of 800 nm in a fundamental

mode transmit through the solid core of a DCF, they suffer from large chromatic

dispersion. Here we shape the refractive index profile of a DCF and place a pair of

inbuilt LPGs inside the fibre core as shown in Fig. 5.1(a) to realize transmit optical

pulses with no chromatic dispersion. The DCF consists of a core, an inner cladding

and an outer cladding section and the refractive index profile of each section is

displayed in Fig. 5.1(b). The coating of the DCF is high refractive index acrylate

coating. Alternatively, a DCF can also be realized by replacing the outer cladding

section with low refractive index coating. Here our design uses the former

approach which is same as the structure of most commercial DCFs.

Chapter 5

89

Fig. 5.1: (a) Schematic structure of a DCF intergrated with a pair of gratings for

achieving zero net chromatic dispersion. (b) Refractive index profile of the DCF.

Chapter 5

90

A pair of long period gratings are used in the conversion from a fundamental mode

to a higher order mode LP02 (Fig. 5.1(a)). The net zero chromatic dispersion to

optical pulses is realized by arranging the dispersion parameters of fundamental

mode pulses and higher order mode pulses into an opposite sign. However a

commercially available DCF has normal dispersion for both a fundamental mode

and a higher order mode. We shaped the refractive index profile of a DCF (Fig.

5.1(b)) to let the higher order mode have anomalous waveguide dispersion

exceeding the material dispersion value, giving a net anomalous dispersion value.

Fig.5.2: Mode evolution for mode LP01 and mode LP02 at wavelengths 725 nm

(blue), 775 nm (green) and 825 nm (red) for the w-DCF. The refractive index

profile is shown in background (black).

Chapter 5

91

Figure 5.2, shows the simulated radial intensity distribution of the LP01 and LP02

modes at different wavelength components of 725 nm, 775 nm and 825 nm using

simulation software Optifibre (Optiwave). As the wavelength component

increases, the fraction of power of the LP02 mode in higher-index regions increases

with the longer wavelength as the mode confines towards the core-center. For a

given wavelength light in higher-index regions travels slower than that in lower-

index regions. The blue components in the LP02 mode travel faster than the red

components, and the LP02 mode has anomalous waveguide dispersion in the fibre.

The core diameter (d1), the thickness of the index dip region outside the core (d2)

and the thickness of the inner cladding region (d3) as shown in Fig. 5.1 affect the

waveguide dispersion of the LP02 mode in the DCF. Figure 5.3 reveals the

simulated waveguide dispersion Dw of the DCF affected by d1, d2, and d3. The

wavelength of the peak waveguide dispersion increases with the increase of d1, d2

and d3.

Chapter 5

92

Fig. 5.3: Waveguide dispersion Dw of a higher order mode in the DCF (a) for

different core diameters d1, (b) for different thicknesses of the index dip region d2,

and (c) for different cladding thicknesses d3.

Chapter 5

93

The core diameter d1, the thickness of the index dip region d2 and the thickness of

the inner cladding region d3 are chosen as 3.κ m, 2.6 m and 14.2 m,

respectively, to obtain the anomalous dispersion peak value at the wavelength of

800nm for the designed core NA. By achieving a high waveguide dispersion value

for an LP02 mode in the fibre, the positive value for the total dispersion at the

designed wavelength of 800 nm can be obtained.

Figure 5.4: (a) Dispersion values for the higher order mode. (b)

Dispersion values for the fundamental mode.

Chapter 5

94

After considering the low index values (that can be achieved through adding

fluorine in silica during fabrication process) in the dip region, we were able to

achieve anomalous waveguide dispersion of +180 ps/nm·km for the wavelength of

800 nm. Figure 5.3 shows the dispersion value for LP02 and LP01.Taking away the

material dispersion -110 ps/nm·km, the total dispersion value for the LP02 mode in

the fibre is ~ +86 ps/nm·km. The waveguide dispersion for the fundamental (LP01)

mode is -197 ps/km·nm, making total dispersion value -324 ps/km·nm. By keeping

the ratio of (L1+L3): L2=1:3.77, the DCF can realize total dispersion of zero to

optical pulses at the central wavelength of 800 nm (Figure 5.5). The dispersion of

mode LP02 of our designed fibre at the wavelength of 800 nm is slightly smaller

than that of a HOM fibre at 770 nm with D=+112.7 ps/(nm·km) [159]. Therefore

our designed fibre requires a longer potion of fibre propagating in mode LP02,

which would lead to the benefit of reducing the fibre nonlinearity since the

nonlinearity of mode LP02 is lower than that of the fundamental mode.

Fig. 5.5: Total dispersion of the fibre under the condition (L1+L3) :L2 = 1:3.77.

Chapter 5

95

In this case, the third order dispersions (TOD) of the fundamental mode and the

LP02 mode are 4.6×10-5

ps3/m and -5.4×10

-4 ps

3/m, respectively. For comparison,

the third order dispersions of a HOM fibre is -0.0002229ps3/m [159] which is

slightly lower than the TOD of the LP02 mode in the DCF. However the third order

dispersion of mode LP01 has an opposite sign to that of mode LP02, thus it can

partially compensate for the third order dispersion of mode LP02 in the DCF. The

net accumulated TODs achieved with a grating pulse compressor or a prism pulse

compressor is ~ +1.1×10-4

ps3/m and ~ -1.3×10

-4 ps

3/m, respectively. The TOD of

the DCF is slightly higher than that from a pair of gratings and a pair of prisms.

However, dispersion compensation using gratings or prisms is bulky and induces

high loss to optical pulses.

The mode effective area of the DCF is found as Aeff=180 µm2, which is 12 folds

larger than that of a HOM fibre [159]. Therefore the DCF induces less nonlinearity

and is more tolerant to the high peak power of optical pulses than HOM fibres. For

optical pulses with a pulse width of 100 fs and a peak power of 1 kW, the

dispersion lengths (Ld) of mode LP01 and mode LP02 are 9 and 34 cm while the

nonlinear length (Lnl) of mode LP01 and mode LP02 are 27 and 232 cm, respectively.

The pulse propagation in the core of the DCF is basically linear and the impact of

nonlinear effects can be negligible.

Chapter 5

96

5.3 Mode conversions

LPGs can induce a single periodic perturbation to a mode in a fibre and perform a

mode conversion from one mode to another mode [167]. Ideally, with an

appropriate set of gratings, any number of modes with arbitrary amplitudes and

phases can be converted into any other modes with arbitrary phase and amplitude

obeying the energy conservation law [167-174]. In an optical fibre, a LPG gives

the efficient mode conversion between two forward propagating modes by

matching the grating period (Λ) to the effective index (neff) difference between two

co-propagating modes at a specific resonant wavelength λres. The period of the LPG

is given by relation [167],

0102 LPLP

res

nn

, (5.1)

where 01LPn and 02LPn are the effective index of the two coupled modes. Here, the

conversion between fundamental and higher order modes (LP01 and LP02) is

performed by a pair of inbuilt LPGs (Fig. 1(a)). For the DCF shown in Fig.1, 01LPn =

1.45526, and 02LPn =1.45515; therefore the period of the LPG Λ is 6.7 mm for the

specific resonance wavelength res=800 nm. The grating refractive index

modulation is chosen to be 0.0005, the maximum grating refractive index

modulation because higher refractive index modulation can induce higher

perturbation to a mode in a fibre and be more effective in mode conversion. The

Chapter 5

97

current LPG fabrication method using a UV laser can have index modulation of

0.0005 for the grating in a silica fibre [175-178]. Using the coupled-mode theory,

the transmission function of the LPG at any wavelength is given by [179],

,

)]1([1

])]1([1[sin

1),(2

22

0

res

resgL

T (5.2)

where T is the transmission function of a LPG filter, Lg is the grating length, is

the coupling constant for the grating,

,),,(4

10201

2

0 dxdyEEzyxnk LPLP (5.3)

),,,(),,(),,( 2

0

22zyxnzyxnzyxn (5.4)

where ),,(2zyxn is the periodic refractive index perturbation of the grating,

),,(2

0 zyxn is the index profile of the waveguide, ),,(2zyxn is the grating index

profile, is the frequency of electric field oscillation, 0 is the permittivity in free

space, 01LPE is the mode field of LP01, and

01LPE the mode field of LP02.

The broad mode conversion bandwidth is achieved through the optimization of the

grating length and the grating period [180-182]. The grating period is calculated by

Equation (5.1) and is determined by the refractive index profile of the DCF.

However the grating period can still be detuned slightly and also keep mode LP02 at

anomalous dispersion.

Chapter 5

98

Fig. 5.6: Mode conversion of the first (a) and the second (b) long

period grating are the intensity of input in mode LP01, input in mode LP02, output in

mode LP01 and output in mode LP02.

Figure 5.6 reveals the transmission spectrum of the two modes by the two gratings

where a refractive index modulation is 0.0005. The mode conversions of the first

and second LPGs are slightly different due to different lengths of the used LPGs.

The length of the first LPG is 16.42 mm while the length of the second LPG is

32.75 mm. Variation of the lengths of two LPGs is due to the optimization of

different transmission peaks of two modes for converting LP01 to LP02 and vice

outoutinin HFHF IIII ,,,

Chapter 5

99

versa. All optical intensity at a wavelength of 800 nm is converted from a

fundamental mode in the core to a higher order mode in the inner cladding by the

first grating and recovered back to a fundamental mode by the second grating as

shown in Fig. 5.6. The two gratings can convert more than 99% of light from a

fundamental mode into a higher order (LP02) mode and back to a fundamental

mode over a bandwidth of 23 nm at the central wavelength of 800 nm. The

bandwidth of 100 fs optical pulses is 15 nm. Therefore the LPGs can sufficiently

convert optical pulses with a pulse width over 100 fs between the fundamental

mode and the LP02 mode.

For LPG fabrication, among different fabrication methods, realization of LPG in

the core of the hydrogen loaded optical fibre by exposing to KrF laser (248nm), is

established method currently in use [93,30,116-119,179-190]. The 2-3% Hydrogen

loading for few hours will increase defects at Ge sites in Germanium doped fibre

core before UV exposure. After laser writing the fibre is annealed to stabilize the

achieved refractive index change by annealing the fibre. This process extract the

hydrogen from the fibre, further stopping hydrogen to increase the defects at Ge

sites that would have raised the average index of fibre resulting the shift in the

resonance peak to longer wavelength. Annealing is generally carried out at 150

degree centigrade for 10 hours. The long period of the gratings is robust to the

bending of the fibre in practical use. The change of the refractive index of grating

period is normally less than 0.1% due to bending, which can be ignored. In

addition, there is a slight change in the resonance peak of the mode conversion for

Chapter 5

100

1-2 % change in length of the grating period due to stretching or bending of fibre,

which can also be ignored.

5.4 High efficient operation of two color light

In many applications including nonlinear endoscopy, an optical fibre is required to

deliver ultra-short optical pulses with a central wavelength of 800 nm as well as to

collect the continuous wave (CW) light with high collection efficiency at other

wavelengths. The design of the DCF, discussed in Figs. 5.1-5.6, can meet those

demands. First, optical pulses are delivered through the core of the DCF. As been

demonstrated in Fig. 5.6, the loss of the DCF to optical pulses, caused by a pair of

inbuilt gratings, is less than 1%. Second, the DCF can also be used for the

backward collection of visible light with a high efficiency. The NA of the core and

the inner cladding of the DCF which are determined by the refractive index

contrast of the core and the trench, and the index contrast of the inner cladding and

the outer cladding, respectively, as shown in Fig. 5.1(b). The NA of the core and

the inner cladding of the DCF are ~ 0.13 and ~ 0.21, respectively. Here, the

wavelength used for calculating the NA of the inner cladding is 521 nm, which is

chosen as the emission wavelength of fluorescein. Figure 5.7(a) shows the

schematic diagram of the DCF for the collection of visible light and dependence of

collection efficiency on the NA of the visible light beam. The collection efficiency

is calculated by calculating the coupling efficiency of the DCF to visible light using

Chapter 5

101

ray optics. The rays with angless less than 120, the corresponding beam angles for

NA of 0.21, can propagate through the DCF through total reflection and the rays

with angless larger than 120 are leaked from the DCF during the propagation and

could not be successfully collected by the DCF. As shown in Fig. 5.7(b), the DCF

can collect 100% of the visible light if the beam NA is less than 0.21. If the optical

pulsed beam from the core of the DCF fills the full aperture of the objective lens,

the NA of the CW beam coupled back by the objective will be equal to the core NA

of the DCF. In this case, the DCF can fully collect all the visible light coupled

through the objective lens.

Fig.5.7 (a) Schematic diagram of the DCF for operating two color light. (b) The

fibre collection efficiency η of the visible light versus the NA of the beam coupled

by an objective.

Chapter 5

102

High NA and large size of the inner cladding can increase the collection of the

visible light. However the size of the inner cladding of the DCF also affects the

dispersion parameter of the LP02 mode (Fig. 5.3). We can maximize the NA and

size of the inner cladding of the DCF without sacrificing anomalous dispersion of

mode LP02.

5.5 Conclusion

In summary, a double-clad solid silica fibre integrated with inbuilt LPGs has been

designed based on the fabrication limitation of the manufacturing process. The

DCF demonstrates two values of the NA for near-infrared and visible beams.

Therefore, it can realize the chromatic dispersion compensation within the optical

fibre at a wavelength of 800 nm as well as provide the high efficiency collection of

visible light through its inner cladding. The loss of the DCF to near-infrared optical

pulses is less than 1%. The designed DCF, which can realize the dispersion

compensation and low loss to the excitation pulses, is important for nonlinear

endoscopy systems to become portable and be able to use a low power and cost

effective femtosecond fibre lasers.

Chapter 6

103

Chapter 6

Conclusion

6.1 Thesis conclusion

Fibre-optic nonlinear imaging is one of the best tools available presently to access

internal organs and tissue surfaces for real time diagnostic imaging and minimally

invasive surgical procedures. The miniaturization of the optical probe and making

whole system more versatile, easy to operate and robust portable has been the

biggest challenge of all times since its beginning. Resolution improvement of these

devices adds their importance in the long term study of the disease development

and the effectiveness of drugs at a sub-cellular level. The low optical power

Chapter 6

104

requirement with highly efficient utilization of excitation beams is important to

make these systems use low cost pulsed femtosecond laser sources. Our

experimental results show that the use of hollow-core photonic crystal fibre (HC-

PCF) in the endomicroscopy system provides options for the broadband excitation

and collection enabling the simultaneous use of multiple fluorescent markers with

different excitation peaks. There is no need of a separate fibre to collect back

fluorescent signal as the cladding of the same HC-PCF collects back fluorescent

signal in the system with the integrated HC-PCF and GRIN lens probe used. The

fibre has a higher nonlinear power threshold than other fibres and the optical power

level required to handle by the fibre is also reduced by the elimination of the

requirement of the use of pre-chirp unit for the chromatic dispersion compensation

that comprises mostly grating pairs which in some cases loses more than 80%

optical power and hinders the device tuneability for different wavelength operation.

Resolution improved fibre-optic nonlinear endoscopy could revolutionize the

clinical research. Super-resolution technology is highly desired features in fibre-

optic based systems in research communities. Towards that direction our study of

the propagation of doughnut beams, to implement STED like super-resolution

imaging features in fibre-optic nonlinear endomicroscopy is valuable. The use of a

single wavelength helps to optimise optics for excitation. Following the successful

demonstration of the two wavelength STED system using a vortex beam in the

fibre-optic endomicroscopy system and the single wavelength bench-top STED in

Chapter 6

105

microscope system we have characterized the propagation of the doughnut beam

for different input polarization states through the HC-PCF fibre at a single

wavelength. Our experimental study shows the way to achieve the doughnut beam

and the propagation by varying the coupling NA for radial and circular polarization

states.

Our new robust solid silica fibre design compensates for the group velocity

dispersion within its designed length by using the mode conversion technique for

femtosecond excitation pulses at the wavelength of 800 nm and also collects the

fluorescent signal through its cladding. The optical loss of less than 1% is achieved

in the mode conversion within the fibre, making it possible to use low power, cost

effective pulsed femtosecond fibre /semiconductor laser in fibre-optic nonlinear

imaging systems.

Chapter 6

106

6.2 Future work

Implementation of a micro-electro mechanical system (MEMS) based probe

scanning system to carry out broadband excitation and collection from biological

samples for multiple fluorescent markers simultaneously will be worth trying.

Further the characterization of newly designed fibre and its use in nonlinear

endoscopy imaging based super-resolution imaging is worth continuing study in

the future. Fibre-optic based single wavelength super-resolution imaging of

biological samples using various long emitting dyes will be the first study of its

kind towards the demonstration of a simple fibre-optic super-resolution imaging

system. Imaging of dyes and bio-samples for different input polarization states with

the HC-PCF fibre is important towards that goal. Detailed study of characteristic of

different dyes particularly long emitting dyes and auto fluorescence for the use in

single wavelength super-resolution imaging is necessary. Pulse characterization

and synchronization using pulse measurement techniques like second harmonic

generation is necessary to confirm the pulse overlap at the focus of the endoscope

probe example that uses a gradient index lens and the scanning mechanism like

using MEMS is necessary to realize miniaturize fibre-optic based super-resolution

endoscopy. Super-resolution imaging study of thick samples using single or

multiple wavelengths is necessary before we can use these techniques for long term

imaging of live suspects.

Bibliography

107

Bibliography

[1] Kobat, D., “Deep tissue multiphoton microscopy using longer wavelength

excitation”, Opt. Express. 17, 13354–13364, (2009)

[2] C.Wang, L.Oiao, Z. Mao, Y.Cheng, and Z.Xu, “Reduced deep-tissue image

degradation in three-dimensional multiphoton microscopy with concentric

two-color two-photon fluorescence excitation”, J. Opt. Soc. Am. B, 25,

976-982, (2008).

[3] D. Brid and M. Gu, “Two-photon fluorescence endoscopy with a micro-

optic scanning head”, Opt. Lett. 28, 1552-1554, (2003).

[4] M. T. Myaing, D.J MacDonald, and X.Li, “Fibre-optic scanning two-

photon fluorescence endoscope” , Opt. Lett. 31, 1076-1078, (2006).

[5] L. Fu and M. Gu, “Fibre-optic nonlinear optical microscopy and

endoscopy” J. Microsc. 226, 195-206, (2007).

Bibliography

108

[6] W.R Zipfel, R. M Williams and W. W. Webb, “Nonlinear magic:

Multiphoton microscopy in the biosciences,” Nat. Biotechno. 21, 1369-

1377, ( 2003).

[7] A. Diaspro, G. Chirico and M. Collini “ Two-photon fluorescence

excitation and related technique in biological microscopy,” Quat. Rev.

Biophys, 38, 2 (2005).

[8] D. Brid and M. Gu, “Resolution improvement in two-photon fluorescence

microscopy with a single-mode fibre”, Appl. Opt. 41, 1852-1856, (2002).

[9] L. Fu, X. Gan, and M. Gu, “Nonlinear optical microscopy based on double-

clad photonic crystal fibres”, Opt. Express. 13, 5528-5534, (2005).

[10] D.Yelin, B. E. Bouma, S.H.Yun, and G.J. Tearney , “ Double-clad fibre for

endoscopy”, Opt. Lett. 29 , 2408-2410, (2004).

[11] S-P Tai, M. C. Chan, T. H Tsai, S-H Guol, L-J Chen, and C-K Sun, “Two-

photon fluorescence microscope with hollow core photonics crystal fibre,”

Opt. Express 12, 6122-6128, (2004).

[12] D. Bird and M. Gu “Compact two-photon fluorescence microscope based

on a single-mode fibre coupler”, Opt. Lett. 27, 1031-1033, (2002).

[13] B. Ha Lee, J. B. Eom, K. S. Park , S. J. Park, and M. J. Ju, “ Specialty

Fibre Coupler: Fabrications and Applications”, J. of the Opt.Soc. of Korea,

14, 326-332, (2010).

Bibliography

109

[14] L. Fu, X. Gan, D. Bird, M. Gu, “ Polarisation characteristics of a 1x 2 fibre

coupler under femtosecond pulsed and continuous wave illumination” , Opt.

and Laser Technol. 37, 494-497, (2004)

[15] L. Fu, X. Gan, and M. Gu “Use of a single-mode fibre coupler for second-

harmonic-generation microscopy”, Opt. Lett. 30, 385-387, (2005).

[16] L. Fu and M. Gu “Double-clad photonic crystal fibre coupler for compact

nonlinear optical microscopy imaging” Opt. Lett. 31, 1471-1473, (2006).

[17] S. Y. Ryu, Hae Y. Choi, M.J.Ju, J. Na, W.J.Choi and B.Ha.Le, “ The

Development of double-clad fibre and double-clad fibre coupler for fibre

based biomedical imaging system”, J. Opt. Soc. Ko. 13 , 310-315 (2009).

[18] H. Bao, S. Y. Ryu, B. H. Lee, W. Tao, and M. Gu, “Nonlinear

endomicroscopy using double-clad fibre coupler ,” Opt. Lett. 35, 995-997,

(2010).

[19] W. Jung, S. Tang, D. T. McCormic, T. Xie, Y-C. Ahn, J.Su,I V. Tomov,

T.B. Krasieva, B. J. Tromberg, and Z. Chen, “Miniaturized probe based on

a microelectromechanical system mirror for multiphoton microscopy,” Opt.

Lett. 33, 1324-1326, (2008).

[20] D. R. Rivera, C. M. Brown, D. G. Ouzounov, I. Pavlova, D. Kobat, W. W.

Webb, and C. Xu, “Compact and flexible raster scanning multiphoton

endoscope capable of imaging unstained tissue,” PNAS, 108, 17598-17603,

(2011).

Bibliography

110

[21] Y. Kremer, J.-F. Léger, R. Lapole, N. Honnorat, Y. Candela, S. Dieudonné,

and L.Bourdieu, “A spatio-temporally compensated acousto-optic scanner

for two-photon microscopy providing large field of view,” Opt. Express 16

,10066-10076, (2008).

[22] S. Kane and J. Squier, “Grating Compensation of Third-Order Material

Dispersion in the Normal Dispersion Regime: Sub- 100-fs Chirped-Pulse

Amplification Using a Fibre Stretcher and Grating-Pair Compressor,”

Quan. Elect. 31, 2052-2057, (1995).

[23] R. Du, R. Jiang and L. Fu, “Enhanced dispersion compensation capability

of angular elements based on beam expansion,” Opt. Express 17, 16415-

16422, (2009).

[24] H. Bao and M. Gu, “ Reduction of self-phase modulation in double-clad

photonic crystal fibre for nonlinear optical endoscopy” , Opt. lett. 34, 148-

150, (2009).

[25] H. Bao, J. Allen, R. Pattie, R. Vance and M. Gu “Fast handheld two-photon

fluorescence microendoscope with a 475 μm - 475 μm field of view for in

vivo imaging” Opt. Lett. 33, 1333-1335, (2008).

[26] C. Xu and W. W. Webb “Measurement of two-photon excitation cross

sections of molecular fluorophores with data from 690 to 1050nm”, Opt.

Express 13, 481-491, (1996).

Bibliography

111

[27] M. A. Albota, C. Xu and W. W. Webb “Two-photon fluorescence

excitation cross sections of bimolecular probes from 690 to 960nm”, Appl.

Opt. 37, 7352-7356, (1998).

[28] Y. R. Shen, The Principles of Nonlinear Optics, John Wiley & Sons, New

York, 1984

[29] X. G. Xu, S. O. Konorov, J. W Hepburn, and V. Milner, “Background free

coherent Raman spectroscopy by detecting the spectral phase of molecular

vibration,” Opt. Lett. 33, 1177-1179, (2008).

[30] U. Utzinger, R. R. Richards-Kortum, “Fibre-optic probes for biomedical

optical spectroscopy”, J. biomed. Opt. 8, 121-147, (2003).

[31] L.Wang, H.Y. Choi, Y. Jung, B.H. Lee, K-T. Kim, “Optical probe based on

double-clad optical fibre for fluorescence spectroscopy”, Opt. Express 15,

17681-17691, (2007).

[32] Y. Wu and X. Li , “Combined influences of chromatic aberration and

scattering in depth-resolved two-photon fluorescence endospectroscopy”, J.

Biomed. Opt. Express 1, 1234-1243, (2010).

[33] F. Helmchen and W. Denk, “Deep tissue two-photon microscopy”, Nature

Meth. 2, 932-940, (2005).

[34] Denk, W., Strickler, J. H. & Webb, W.W. “ Two-photon laser scanning

fluorescence microscopy”. Science 248, 73-76, (1990).

Bibliography

112

[35] J. M Girkin and D. Wokosin, Practical multi-photon microscopy In

confocal and Two-photon microscopy : Foundation, Applications and

Advances (Wiley-Liss & sons Inc., 2001).

[36] M. Gu and L. Fu “Three-dimensional image formation in fibre-optical

second-harmonic-generation microscopy”, Opt. Express 14, 1175-1181,

(2006).

[37] L. Fu, A. Jain, H. Xie, C. Cranfield and M. Gu “ Three-dimensional

nonlinear optical endoscopy” , SPIE 4, 400501 (1-3), ( 2007 ).

[38] M. Gu, “Principles of Three dimensional imaging in confocal

microscopes”, world scientific publishing, 1-274, (1996).

[39] G. Chirco, F. Cannone, Baldini G and Diaspro A “ A two-photon thermal

bleaching of single fluorescent molecules” , J. Biophys. J. 84, 588-598,

(2003b).

[40] W-F. Cheong, S.A.Prahl, and A.J.Welch, “ A Review of the optical

properties of Biological Tissue”, J. Quantum. Electron. 26 , (1990).

[41] R.K Wang, “Modeling optical properties of soft tissue by fractal

distribution of scatterers”, J. Modern Opt., 47 ,103-120, (1999).

[42] C.Mar Blanca and C. Saloma, “Monte Carlo analysis of two-photon

fluorescence imaging through a scattering medium”, Appl. Opt., 37, 8092-

8102, (1998).

Bibliography

113

[43] X. Deng, X. Gan and M. Gu, “Monte carlo simulation of multiphoton

fluorescescence microscopic imaging through inhomogeneous tissuelike

turbid media”, J. Biomed. Opt. 8 ,440-449, (2003).

[44] H.C. Gerristsen, C.J. DeGrauw, “Imaging of optically thick specimen using

two-photon excitation microscopy” Microsc. Res Tech. 47, 206-209,

(1999).

[45] J. W. McLean, J. D Freeman, R. E. Walker, “Beam spread function with

time dispersion”, Appl. Opt. 37, 4701, (1998 ).

[46] G.A Blab, P.H Lommerse , L, Cognet, G. S Harms and T. Schmidt “ Two-

photon excitation action cross-section of the auto fluorescent proteins”

Chem. Phys. Lett. 350, 71-77, (2001).

[47] E. Abbe, Gesammelte abhandlugen ( G. Fischer Jena, 1904)

[48] Q. Li, S. S.H. Wu, K.C.Chou, “Subdiffraction-limit two-photon

fluorescence microscopy for GFP-tagged cell imaging”, Biophys. J. , Vol.

97, 3224-3228, (2009).

[49] T. A. Klar and S. W. Hell, “subdiffraction resolution in far-field

fluorescence microscopy”. Opt. Lett. 24, 954-956, (1999).

[50] P.Dedecker, C.Flors, J-I Hotta, H. Uji-I and J. Hofkens “3D

Nanoscopy:bringing biological nanostructures into sharp focus”,

Angew.Chem Int. 46,8330-8332 ,(2007).

Bibliography

114

[51] E.Betzig, G.H. Patterson, R. Sougrat, O.W. Lindwasser, S. Olenych, J. S

Bonifacino, M. W. Davidson, J. Lippincott-Schwartz, H.F. Hess, “Imaging

intracellular fluorescent proteins at nanometer resolution”,Science

313,1642-1645, (2006).

[52] P.Dedecker, J. Hofkens, and J-I Hotta, “Diffraction-unlimited optical

microscopy”, Mater. Today, 11, 12-21, (2008).

[53] E. Haustein and P. Schwille, “Trends in fluorescence imaging and related

techniques to unravel biological information”, HFSP J., 1(3), 169-180,

(2007).

[54] M. Gu , T. Tannous, and C. J. R. Sheppard, “ Theree dimensional confocal

fluorescence imaging under ultrashort pulse illumination” Opt. commun.

117, 406-412, (1995).

[55] G.cox and C. J.R. Sheppard, “Practical limits of Resolution in confocal and

Non-linear Microscopy”,Microsc. Res. and Tech. 63, 18-22 (2004).

[56] G. donnert et. al ,“Macromolecular-scale resolution in biological

fluorescence microscopy”, PNAS, 103, 11440- 11445, (2006).

[57a] B.Huang, M.Bates, and X.Zhuang, “Super-Resolution Fluorescence

Microscopy”, Annual Rev. Biochem. 78, 993-1016, (2009).

[57b] B. R. Masters, P. T. C. So, “Confocal microscopy and multi-photon

excitation microscopy of human skin in vivo”, Opt. Express 8, 2-10,

(2001).

Bibliography

115

[58a] D.Wildanger, E.Rittweger, L. Kastrup and S.W.Hell, “STED microscopy

with a supercontinuum laser source”, Opt. Express 16, 9614-9621, (2008).

[58b] W. Denk, K. Svoboda, “Photon Upmanship : Why Multiphoton imaging is

more than a Gimmick”, Neuron 18, 351-357, (1997).

[59] V. Westphal and S. W. Hell, “Nanoscale resolution in the focal plane of an

optical microscope, “Phys. Rev. Lett. 94, 143903, (2005).

[60] B.P, Cormack, R.H. Valdivia and S. Falkow, “FACS-optimized mutants of

the green fluorescent protein (GFP)”, Gene, 173, 33-38, (1996).

[61] K. I Willing, S.O. Rizzoli, VWestphal, R. Jahn, and S.W Hell, “STED

microscopy reveals that synaptotagmin remains cluster after synaptic

vesicle exocytosis” , Nature. 400,935-939, (2006).

[62] G. Vicidomini, G. Moneron, C.Eggeling, E. Rittweger, and S.W Hell,

“STED with wavelengths closer to the emission maximum”, Opt. Express.

20, 5225-5236, (2012).

[63] J. R Moffitt, C. Osseforth, and J. Michaelis, “Time-gating imporves the

spatial resolution of the STED microscopy,” Opt. Express 19, 4242-4254,

(2011).

[64] G. Vicidomini, G. Moneron, K.Y. Han, V. Westphal, H. Ta, M. Reuss, J

Engelhardt, C. Eggeling, and S. W. Hell, “ Sharper low-power STED

nanoscopy by time gating, : Nat. Methods 8, 571-573, (2011).

Bibliography

116

[65] S.T Hess, T.P.K Girirajan, M. D.Mason, “Ultra-High Resolution Imaging

by Fluorescence Photoactivation Localization Microscopy” Biophys J., 91,

4258-4272, (2006).

[66] M.J.R. Rust, M.B. Bates, X.Z.Zhuang, “Sub-diffraction-limit imaging by

stochastic optical reconstruction icroscopy (STORM)”,Nat. Methods, 3,

793-796, (2006).

[67] M. Aeschlimann,m.bauer, D.Bayer, T.Brixner, F.J. G.d.Abajo,W.Pfeiffer,

M.Rohmer, C.spindler and F.Steeb, “Adaptive subwavelength control of

nano-optical fields”, Nature 446, 301-304, (2007).

[68] R.Menon, P Rogge and H-Y Tsai, “Design of diffractive lenses that

generate optical nulls without phase singularities”, Opt. Soc. Am. 26, 297-

304, (2009).

[69] D.Wildanger, R. Medda, L.Kastrup and S.W.Hell, “A compact STED

microscope providing 3D nanoscale resolution”, J. Microsc. 236, 33-43,

(2009).

[70] Baer, S., “Method and apparatus for improving resolution in scanned

optical system,” U.S. Patent #5,866,911,(filed July 15, 1994, published Feb.

29, 1996 as International Application WO 9606369A2, issued Feb. 2, 1999)

[71] S. W. Hell and J. Wichmann, “Breaking the diffraction resolution limit by

stimulated emission: Stimulated emission - depletion fluorescence

microscopy”, Opt. Lett. 19, 780-782, (1994).

Bibliography

117

[72] C. Eggeling, C. Ringemann, R. Medda, G.Schwarzmann, K. Sandhoff, S.

Polyakova, V.N. Belov, B. Hein, C. Von. Viddendorff, A. Schonle, and S.

W. Hell, “ Direct observation of the nanoscale dynamics of membrance

lipids in a living cell,” Nature 457 ,1159-1162, (2009).

[73] J. B. Ding, K. T. Takasaki and B.L. Sabatini, “Supraresolution imaging in

brain slices using stimulated-emission depletion two-photon laser scanning

microscopy,” Meirpm 63,429-437 (2009).

[74] T. Grotjohann, I. Testa, M. Leutenegger, H. Bock, N. T. Urban, F. Lavoie-

Cardinal, K. I. Willing, C. Eggeling, S. Jakobs, and S. W. Hell, “

Diffraction-unlimited all-optical imaging and writing with a photochromic

GFP,” Nature 478,204-208, (2011).

[75] J.Huisken, j.Swoger, F.D Bene, J.Wittbrodt, E .H.K.Stelzer, “Optical

sectioning deep inside live embroys by selective plane illumination

microscopy”,Science, 305, 1007-1009, (2004).

[76] G.Cristobal, E.gil, F.sroubek, J.Flusser, D.Miravet, F.B.Rodriguez,

“Superresolution imaging: a survey of current techniques”, Proc. of SPIE

7074, 0C-18, (2008).

[77] Ilan Davis, “ The ‘Super Resolution’ revolution”, Biochem. Soc. Transac.,

37, part 5 (2009).

[78] Gustaffsson. MGL ,“Nonlinear sturctured-illumination microscopy: Wide-

field fluorescence imaging with theoretically unlimited resolution”, PNAS

102, 13081-13086, (2005).

Bibliography

118

[79] G.Zhu, J.V.Howe, M.Durst, W.Zipfel and C.Zu, “Simultaneous spatial and

temporal focussing of femtosecond pulses “, Opt. Express, 13, 2153-2159,

(2005)

[80] D.M.D Juana, J.E.Oti,Vidal F.Canales and M.P.Cagigal, “Transverse or

axial superresolution in a 4Pi-confocal microscope by phase-only filters”,

Opt. Soc. Am. 20 , 2172-2178, (2003).

[81] P.E.Hanninen, S.W. Hell, J. Salo, E.Soini, C.Cremer, “Two-photon

excitation 4Pi confocal microscope: Enhanced axial resolution microscope

for biological research”, Appl. Phys. Lett. 66, 1698-1700, (1995)

[82] Hell, S.W “Towards fluorescence nanoscopy”, Nature Biotechnology, 21,

1347-1355 (2003).

[83] M. Fernandex-suarez and A. Y. Ting, “Fluorescent probes for super-

resolution imaging in living cells”, Nature Rev. Molec. cell Bio. 9 , 929-

9439, (2008).

[84] S. Zhuo, J. Yan, G. Chen, J. Chen, Y. Liu, J. Lu, X. Zhu, X. Jiang and

S.Xie,“Label-free monitoring of colonic cancer progression using

multiphoton microscopy” J. of Biomed. Opt. Express 2, 615-619, (2011).

[85] W. Zheng. , W. Lau, C. Cheng, KC Soo, M Olivo , “ Optimal excitation-

emission wavelength for autofluorescence diagnosis of bladder tumors” ,

Internat. J. Cancer 104, 477-481, (2003).

[86] M.Z. Lin, M. R. McKewon, Ng HL, T. A. Aguilera, N.C. Shaner, R.E.

Campbell, S.R Adams, L.A Gross, W Ma, T Alber , R.Y Tsien , “

http://librarysystem.swin.edu.au/primo_library/libweb/action/search.do?vl\(freeText0\)=Lau%2c+W&vl\(2608403UI0\)=creator&vl\(2608402UI1\)=all_items&fn=search&tab=remote&mode=Basic&vid=SWIN1&scp.scps=000026491

http://librarysystem.swin.edu.au/primo_library/libweb/action/search.do?vl\(freeText0\)=+Cheng%2c+C&vl\(2608403UI0\)=creator&vl\(2608402UI1\)=all_items&fn=search&tab=remote&mode=Basic&vid=SWIN1&scp.scps=000026491

http://librarysystem.swin.edu.au/primo_library/libweb/action/search.do?vl\(freeText0\)=+Soo%2c+KC&vl\(2608403UI0\)=creator&vl\(2608402UI1\)=all_items&fn=search&tab=remote&mode=Basic&vid=SWIN1&scp.scps=000026491

http://librarysystem.swin.edu.au/primo_library/libweb/action/search.do?vl\(freeText0\)=+Olivo%2c+M&vl\(2608403UI0\)=creator&vl\(2608402UI1\)=all_items&fn=search&tab=remote&mode=Basic&vid=SWIN1&scp.scps=000026491

Bibliography

119

Autofluorescent Proteins with excitation in the optical window for intravital

imaging in Mammals”, Chem. and Biol. 16, 1169-1179, (2009).

[87] P. Bianchinia, B. Harkea, S. Galiania, G. Vicidominia, and A. Diaspro

“Single-wavelength two-photon excitation–stimulated emission depletion

(SW2PE-STED) superresolution imaging”, PNAS 109, 6390–6393, (2012).

[88] T. scheul, C. D. Amico, I. Wang and J-C. Vial, “Two-photon excitation and

stimulated emission depletion by a single wavelength”, Opt. Exp. 19,

18036-18041, (2011).

[89] D. Baddeley, D. Crossman, S. Rossberger, J. E. cheyne, J. M. Montgomery,

I. D Jayasinghe, C. Cremer, M. B Cannell, C. Soeller “ 4D Super-resolution

microscopy with conventional fluorophores and single wavelength

excitation in optically thick cells and tissues”, PLoS ONE, 6, 1-10, (2011).

[90] I.Iiev , R.Waynant, I.Gannot, A.Gandjbakhche, “Simple fibre-optic

confocal microscopy with nanoscale depth resolution beyond the diffraction

barrier”, Rev. Scient. Inst. 78, 093703, (2007).

[91] Y.Wu,Y.Leng, J.Xi and Z.Li, “Scanning all-fibre-optic endomicroscopy

system for 3D nonlinear optical imaging of biological tissues”, Opt.

Express 17, 7907 – 7915, (2009).

[92] D. Bird and M. Gu “Fibre-optic two-photon scanning fluorescence

microscopy”, J. Microsc. 208, 35-48, (2002).

Bibliography

120

[91] J. C. Knight, J. Arriaga, T.A. Briks, O. Blanch, W.J. Wadsworth and P.St. J.

Russell “ Anomalous Dispersion in Photonic Crystal Fibre”, IEEE Phot.

Technol. Lett. 12, 807-809, (2000).

[92] Y-Z Xu, X-M.Ren, Z.Xia, Y-Q Huang, “ A fully vectroial effective index

method for Accurate dispersion calculation of photonic crystal fibres”,

China Phys. Lett. 23 (2006).

[93] W.Jung, S. Tang, T. Xie, D. T. McCormick, Y.Ahn, J.Su, I. V. Tomov.

T.B.Krasieva, B. J.Tromberg, Z. Chen “Miniaturized probe using 2 axis

MEMS scanner for endoscopic multiphoton excitation microscopy”, Proc.

of SPIE 6851, 0D-7, (2008).

[94] G.A. Stanciu, C. Stoichita, S.G Stanciu, “Scanning laser microscopy : From

Far field to Near Field”, IEEE, ICTON (2009).

[95] B. A Flusberg, E. D Cocker, W. Piyawattanemetha, J. C Jung, E. L M

Chung and M. J Schnitzer , “Fibre-optic fluorescence imaging”, Nature

Meth. 12, 941-949, (2005).

[96] J. C. Knight, J. Arriaga, T.A. Briks, O. Blanch, W.J. Wadsworth and P.St. J.

Russell “Anomalous Dispersion in Photonic Crystal Fibre”, IEEE Phot.

Technol. Lett. 12, 807-809, (2000).

[97] Y-Z Xu, X-M.Ren, Z.Xia, Y-Q Huang, “ A fully vectroial effective index

method for Accurate dispersion calculation of photonic crystal fibres”,

China Phys. Lett. 23 (2006).

Bibliography

121

[98] H. Bao and M. Gu “A 0.4 mm diameter probe for nonlinear optical

imaging”, Opt. Express 17, 10098-10104, (2008).

[99] L. Fu, A. Jain, H. Xie, C. Cranfield and M. Gu “ Nonlinear optical

endoscopy based on a double-clad photonic crystal fibre and a MEMS

mirror” ,Opt. Express 14, 1027-1032, (2006).

[100] H. Kano and H. Hamaguchi, “In vivo multi-nonlinear optical imaging of a

living cell using a super continuum light source generated from a photonic

crystal fibre” Opt. Express 14, 2798-2804, ( 2006).

[101] L. Fu and M. Gu “Polarization states Anisotropy in Fibre-optic second

harmonic generation microscopy” Opt. Express 16, 5000-5006, (2008).

[102] R. L. Harzic, M. Weinigel, I. Riemann, K. König, and B. Messerschmidt,

“Nonlinear optical endoscopy base on a compact two axis piezo scanner

and a miniature objective lens,” Opt. Express 16, 20588-20596, (2008).

[103] C. L. Hoy, O. Ferhanoğlu, M. Yildirim,W. Piyawattanametha, H. Ra, O.

Solgaard, and A. Ben-Yakar, “Optical design and imaging performance

testing of a 9.6-mm diameter femtosecond laser microsurgery probe”, /

Opt. Express 19 , 10536-10552, (2011).

[104] G.P Agrawal, Nonlinear fibre-optics, Academic Press, (2007)

[105] K. Thyagarajan and B.P.Pal , “Modeling dispersion in optical fibres:

applications to dispersion tailoring and dispersion compensation”, J. Opt.

fibre. Commun. Rep.4, 173-213, (2007).

Bibliography

122

[106] J.Kim, D.K, “An efficient dispersion calculation method for axially

symmetric optical fibres”, Fibre and Integ. Opt., 21,13-29, (2002).

[107] W.A.gambling, H.Matsumura and C.M.Ragdale, “Mode dispersion,

material dispersion and profile dispersion in graded-index single-mode

fibres”, Microwaves, Opt. and Acoust., 3, (1979).

[108] Y. Painchaud, C.Paquet and M.Guy, “FBGs enhance dispersion

compensation”, Laser Focus World, 145-147, (2007).

[109] S. Ramachandran, B. Mikkelsen, L.C. Cowsar M. F. Yan, G. Raybon, L.

Boivin, M. Fishteyn, W.A. Reed, P.Wisk, D. Brownlow, R. G Huff and L.

Gruner-Nielsen , “ All-Fibre Grating-Based Hingher Order Mode

Dispersion Compensator for Broad-Band Compensation and 1000-km

Transmission at 40Gb/s” IEEE Phot. Technol. Lett. 13, 632-634, (2001).

[110] W. Yi, W. Ke-jia, W. Xiu-min, Z. Shuai, “The Heterodyne Research on

Transmission Signal of Three-layered Cascaded Long-Period Fibre

Gratings”, Proc. of SPIE 7381, 1R-10, (2009).

[111] X. Chu, S Zhao, Z. Wu, L. Shi, S Zhan, Y. Li, L Ma, “ New method for side

pumping of double-clad fibre lasers using planar grating written in the inner

cladding”, Proc. of SPIE 7382, 0X-7, (2009).

[112] H.Ke, K.S. Chiang, and J.H.Peng, “Analysis of Phase-shifted Long-Period

Fibre Gratings”, IEEE Photon. Technol. Lett. 10, 1596-1598, (1998).

Bibliography

123

[113] M. Hany, M.H.aly and M.Nassr, “Dispersion compensation in Transmission

using AFBG Chirped with Two Beam Interference Fringe Spacing”, J. App.

Sci. Res. 5, 1622-1629, (2009).

[114] B. Y. Kim, J. N. Blake, H. E, Engan, and H. J. Shaw, “All-fibre acousto-

optic frequency shifter”, Opt. Lett 11, 389-391, (1986).

[115] R. C. Youngquist, J. L. Brooks, and H. J. Shaw, “ Birefringence-fibre

polarization states coupler,” Opt. Lett. 8, 656-659, (1983).

[116] Bashir A. T, Jalil A. , Rosly A. R. “ Fabrication of fibre grating by phase

mask and it’s sensing application” , J. Opto.Elec. and Adv. Mat. 8, 1064-

1609, (2006).

[117] Yuki K. ,Kentaro N. and Tsuneo. M , “Fabrication of long-period fibre

gratings by focused irradiation of infrared femtosecond laser pulses”, Opt.

Lett. 24, 646-648, (1999).

[118] Y. Kondo, K. Nouchi, T. Mitsuyu, M. Watanabe, P. Kazansky, and K.

Hirao, “ Fabrication of long period fibre grating by focused irradiation of

infrared femtosecond laser pulses”, Opt. Lett. 24, 646-648, (1999).

[119] P.G. Kryukov, Yu.V. Larionov, A. A. Rybaltovskii, K.A.Aagorul’ko, A.

Dragomir, D.N. Nikogosyan, A.A. Ruth, “Long-period fibre grating

fabrication with femtosecond pulse ration at different wavelength”, Micro-

electron. Eng. 69 , 248-255 (2003).

[120] T. Erdogan, “Cladding-mode resonances in short- and long-period fibre

grating filters”, J. Opt. Soc. Am. 14, 1760-1773, (1997).

Bibliography

124

[121] J. P. Ying, F. Liu, and R.R. Alfano, “Spatial distribution of two-photon-

excited fluorescence in scattering media: erratum” Appl. Opt. 38, 2151-

2156, (1999).

[122] P. Theer, M. T. Hasan and W. Denk “ Two-photon imaging to a depth of

1000 μm in living brains by use of a 32: OAlTi regenerative amplifier”

,Opt. Lett. 28, 1022-1024, (2003).

[123] M. Oheim , E. Beaurepaire, E. Ghaigneau, J. Mertz and S. Charpak “Two-

photon microscopy in brain tissue: parameters influencing the imaging

depth”, J. Neurosci. Meth. 111, 29-37 (2001).

[124] E. Beaurepaire, M Oheim, J. Mertz “Ultra-deep two-photon fluorescence

excitation in turbid media” , Op. Comm. 188, 25-29, (2001).

[125] C.Wang, L.Oiao, Z. Mao, Y.Cheng, and Z.Xu, “Reduced deep-tissue image

degradation in three-dimensional multiphoton microscopy with concentric

two-color two-photon fluorescence excitation”, J. Opt. Soc. Am., 25,

(2008).

[126] Pang, S., Yeh, A. T., Wang, C. Meissner, K. E. “Beyond the 1/Tp limit:

two-photon excited fluorescence using pulses as short as sub-10-fs” J.

Biomed. Opt. 14, 054041, (2009).

[127] Tang, S. Krasieva, T. B., Chen Z., Tempea, G., Tromberg, B. J. “Effect of

pulse duration on two-photon excited fluorescence and second harmonic

generation in nonlinear optical microscopy” J. Biomed. Opt. 11, 020501,

(2006).

Bibliography

125

[128] Larson, A. M. and Yeh, A. T. “Ex-Vivo characterization of sub-10-fs

pulses”, Opt. Lett. 31, 1681–1683, (2006).

[129] D.Wildanger, J.Buckers, V.Westphal, S.W.Hell and L.Kastrup, “A STED

microscope aligned by design” , Opt. Express, 17, 16100-16110, (2009).

[130] D. Kim, H. Choi, S. Yazdanfar, and P. T. C. So, “Ultrafast optical pulse

deliver with fibres for nonlinear microscopy,” Micros. Res. Tech. 71, 887-

896, (2008).

[131] D. G. Ouzounov, K. D. Moll, M. A. Foster, W. R. Zipfel, W. W. Webb, and

A. L. Gaeta, “Delivery of nanojoule femtosecond pulses through large-core

microstructured fibres,” Opt. Lett. 27, 1513-1515, (2002).

[132] Shephard, J.D., Macpherson, W.N., Maier, R.R.J., Jones, J.D.C.,

Hand,D.P., Mohebbi, D.P., George, A.K., Roberts, P.J., and Knight,

J.C.:‘Single-mode mid-IR guidance in a hollow-core photonic crystal

fibre”,Opt. Express 13, 7139–7144, (2005)

[133] P. J Mosley, W. C. Huang, M. G. Welch, B. J. Mangan, W. J. Wadsworth

and J. C. Knight, “Ultrashort pulse compression and delivery in a hollow-

core photonic crystal fibre at 540 nm wavelength”, Opt. Lett. 35, 3589-

3591, (2010).

[134] F. Benabid, “Hollow-core photonic bandgap fibre: new light guidance for

new science and technology,” Philos Transact A Math Phys Eng Sci., 364,

3439-3462, (2006).

Bibliography

126

[135] J. C. Knight, F. Gérôme, and W. J. Wadsworth, “Hollow-core photonic

crystal fibres for delivery and compression of ultrashort optical pulses,”

Opt. Quant. Elect. 39 (12-13), 1047-1056, (2007).

[136] T. Le, M. Hofer, Z. Cheng, and A. Stingl, J. Darmo, D.P. Kelly, and K.

Unterrainer,“Advances in Fibre Delivery of Ultrashort Pulses at 800 nm”,

Proc. of SPIE 6871, 68712C, (2008)

[137] D. H. Kim, J. U. Kang, and I. K. Ilev, “Advance confocal microscope using

single hollow core photonic bandgap fibre design,” Elect. Lett., 43, 608-

609, (2007).

[138] W. Göbel, A. Nimmerjahn, and F. Helmchen, "Distortion-free delivery of

nanojoule femtosecond pulses from aTi:sapphire laser through a hollow-

core photonic crystal fibre," Opt. Lett. 29, 1285-1287, (2004).

[139] C. L Hoy, N. J Durr, P. Y Chen, W. Piyawattanametha, H. Ra, O. Solgaard,

and A. Ben-Yakar, “Miniaturized probe for femtosecond laser microsurgery

and two-photon Imaging,” Opt. Express 16 , 9996-10005, (2008).

[140] Y. C. Wu, J. F. Xi, M. J. Cobb, and X. D. Li, “Scanning fibre-optic

nonlinear endomicroscopy with miniature aspherical compound lens and

multimode fibre collector,” Opt. Lett. 34, 953-955, (2009).

[141] C. J. Engelbrecht, R. S. Johnston, E. J. Seibel, and F. Helmchen, “Ultra-

compact fibre-optic two-photon microscope for functional fluorescence

imaging in vivo,” Opt. Express 16, 5556-5564, (2008).

Bibliography

127

[142] S. K.Vengalathunadakal,V.M. Murukeshan, S. Padmanabhan, P.

Padmanabhan,“Photonics crystal fibre based dual modality probe for

simultaneous sensing and imaging application”, Opt. Eng. 48, 103601-

103604, (2009)

[143] R.S.Pillai, C.Boudoux, g.Labroille, N.Olivier, I.Veilleux, E.Farge, M.Joffre,

and E.Beaurepaire, “Multiplexed two-photon microscopy of dynamic

biological sampless with shaped broadband pulses”, Opt. Express 17,

12741-12752, (2009).

[144] S. W. Hell and J. Wichmann, “Breaking the diffraction resolution limit by

stimulated emission depletion fluorescence microscopy ”, Opt. Lett. 19,

780-782, (1994).

[145] Willig, K. I., Harke, B., Medda, R. and Hell, S. W., “STED microscopy

with continuous wave beams,” Nat.Methods 4, 915-918, (2007).

[146] J. B. Ding, K. T. Takasaki and B.L. Sabatini, “Supraresolution imaging in

brain slices using stimulated-emission depletion two-photon laser scanning

microscopy,” Meirpm 63, 429-437, (2009).

[147] Baer, S. C., “Forming light beams and patterns with zero intensity central

points,” US Patent application 20100315709, (priority filing date 2007)

(2007).

Bibliography

128

[148] Morozova, K. S., Piatkevich, K. D., Gould, T. J., Zhang, J., Bewersdorf, J.

and Verkhusha, V. V., “Far-RedFluorescent Protein Excitable with Red

Lasers for Flow Cytometry and Superresolution STED

Nanoscopy,”Biophys. J. 99, L13–L15, (2010).

[149] Drobizhev, M., Tillo, S., Makarov, N. S., Hughes, T. E., and A. Rebane, A.,

“Absolute Two-Photon Absorption Spectra and Two-Photon Brightness of

Orange and Red Fluorescent Proteins,” J. Phys. Chem. B. 113, 855–859,

(2009).

[150] Leutenegger, M., Eggeling, C. and Hell, S. W., “Analytical description of

STED microscopy performance,” Opt.Express 18, 26417-26429, (2010).

[151] Min Gu, Hong Kang & Xiangping Li , “Breaking the diffraction-limited

resolution barrier in fibre-optical two-photon fluorescence endoscopy by an

azimuthally-polarized beam”, Sci. Rep. 4, 3627-3637, ( 2014).

[152] Baer, S., “Single wavelength stimulated emission depletion microscopy,”

U. S. Patent #7,816,654, (priority filing date Jan. 16, 2005, published July

27, 2006 as International Application WO2006078857, issued Oct.

19,2010) (2005).

[153] A. Small, “Model of bleaching and acquisition for superresolution

microscopy controlled by a single wavelength”, Biomed. Opt. Express 2,

2934-2949, (2011).

[154] S. C. Baer, “Single wavelength STED microscope”, Proc. of SPIE , Vol.

7903, doi:10.1117/12.875005, 1-8, (2011).

Bibliography

129

[155] Z.Xiaoping, T.Zhihong, “Analysis of transmission characteristics of doubly

clad fibres with an inner cladding made of uniaxial crystal materials”, Opt.

Commun. 204, 127-136, (2002).

[156] L. Fu, A. Jain,C.Cranfield, H.Xie and M. Gu, “Three-dimensional nonlinear

optical endoscopy,” J. Biomed. Opt. 12, 040501-040509 (2007).

[157] L. Fu, X. Gan and M. Gu “ Characterization of gradient-index lens-fibre

spacing toward applications in two-photon fluorescence endoscopy” , Appl.

Opt. 44, 7270-7274, (2005).

[158] Chi Wang, Youxin Mao, Chen Fang,Zhi Tang, Yingjie Yu, Bo Qi,

“Analytical method for designing gradient-index fibre probes”, Opt. Engg. ,

50, 094202, (2011)

[159] K. J. Jespersen, T. Le, L. G. Nielsen, D. Jackobsen, M. E. V. Pederesen, M.

B. Smedemand, S. R. Keiding, B. Palsdottir, “A higher-order-mode fibre

delivery for Ti:Sapphire femtosecond lasers” Opt. Express 18, 7798-7806,

(2010)

[160] S. Ramachandran, M. F. Yan, Jasapara, P. wisk, S. Ghalmi, E. Monberg,

and F. V. Dimarcello, “High-energy (nanojoule) femtosecond pulse

delivery with record dispersion higher-order mode fibre”,Opt. Lett. 30,

3225-3227, (2005)

[161] R.Siddhartha, “Dispersion-Tailored Few- Mode Fibres: A Versatile

Platform for In-Fibre Photonic Devices”, J. lightwave tech., 23, 3426-3443,

(2005).

Bibliography

130

[162] J. W. Nicholson, S. Ramachandran and S. Ghalmi “ A Passively-

modelocked, Yb-doped, figure-eight, fibre laser utilixing anomalous-

dispersion highter-order-mode fibre” Opt. Express 15, 6623-6628, (2007).

[163] S. Ramachandran, S. Ghalmi, J.W. Nicholson, M. F. Yan, P. Wisk, E.

Monberg and F. V. Dimarcello “ Anomalous dispersion in a solid silica

based fibre” , Opt. Lett. 31, 2532-2534, (2006 ).

[164] M. Schultz, O. Prochnow, A. Ruehl, D. Wandt, D. Kracht, S.

Ramachandran and S. Ghalmi “ Sub-60-fs yetterbium-doped fibre laser

with a fibre based dispersion compensation”, Opt. Lett. 32, 2372-2374,

(2007).

[165] S. Ramachandran, J.W.Nicholson, S. Ghalmi, M.F. Yan, P. Wisk, E.

Monberg and F. V. Dimarcello, “ Light propagation with ultralarge modal

areas in optical fibres” , Opt. Lett. 31, 1797-1799, (2006).

[166] P Hlubina, “Group velocity dispersion in fused-silica samples measured

using white-light interferometry with the equalization wavelength

determination”,Optik 113, 149-152, (2002).

[167] C-X Shi, T Okoshi, “Analysis of fibre-optic LP01-LP02 mode converter,”

Opt. Lett. 17, 719-721, (1992).

[168] M. Sumetsky and S. Ramachandran, “Conversion and focusing of optical

fibre modes with superimposed long period gratings”,OThV3,

OFC/NFOEC, (2008).

Bibliography

131

[169] H.Shu and M.Bass, “Calculating the guided modes in optical fibres and

waveguides”,J. Lightwave Tech. 25, (2007).

[170] M-Y chen, J Zhou, “Mode converter based on mode coupling in an

asymmetric dual-core photonic crystal fibre”, J. Appl. Opt. 10, (2008).

[171] A.C Turner, C.manolatou, B S Schimdt, M. Lipson, M.A Foster, J. E

Sharping, A.L. Gaeta “Tailored anomalous group-velocity dispersion in

silicon channel waveguides”, Opt. Express 14, 4357-4362, (2006).

[172] D. Gloge, “Dispersion in weakly guiding fibres”, Appl. Opt. 10, 2442-2445,

(1971).

[173] J. Yao, L.Zhan, Zh. Luo, and Y. Xia, “Modeling and Numerical

Demonstration of Femtosecond-pulse Distortion-Free Delivery Using a

Three-Layered Few-Mode Fibre”, Laser Phys., 18, 455-459, (2008).

[174] S. Ramachandran, J.W.Nicholson, S. Ghalmi, M.F. Yan, P. Wisk, E.

Monberg and F. V. Dimarcello, “ Light propagation with ultralarge modal

areas in optical fibres” , Opt. Lett. 31, 1797-1799, (2006).

[175] H. J. Patric, C. G. Askins, R. W. Mcelhanon, and E. J. friebele, “ Amplitude

mask patterned on an excimer laser mirror for high intensity writing of long

period fibre gratings”, Electron. Lett. 33, 1167-1168, (1997).

[176] Oleg V. Ivanov, “Fabrication of long-period fibre gratings by twisting a

standard single-mode fibre”, Opt. Lett., 30 , 3290-3292, (2005).

[177] A. Cusano, D. Paladino, A.Cutolo, I.D,Villar, I.R.Matias and F. J. Arregui,

“Spectral Characteristics in long-period fibre gratings with nonuniform

Bibliography

132

symmetrically ring shaped coating”, Appl. Phys. Lett. 90, 14110-14115,

(2007).

[178] L. Li, W.Fan, S. Fu, Y. Liu, S. Yuan, X.Dong, “Fabrication and

characteristics of double-clad fibre Bragg Gratings”, Front. Optoelectron.

China 2 , 68-75, (2009).

[179] H.Chen, “Design and simulation of broadband IR integrated waveguide

amplifiers with long-period waveguide grating filter”,J Infrared Mili Terahz

Waves 31, 24-30, (2010).

[180] T.W. MacDougall, S.Pilevar, C. W. Haggans, and M.A.Jackson,

“Generalized Expression for the growth of long period gratings”, IEEE

photon. Technol. Lett. 10, 1449-1451, (1998).

[181] C-L.Lee and P.Han, “Optimal design of single resonant and ultra

broadband long-period fibre grating filters”,Opt. Eng. Lett., 48, 080501-3,

(2009).

[182] H.Chen, “Design and simulation of Broadband IR integrated waveguide

amplifiers with long-period waveguide grating filter”,J Infrared Mili Terahz

Waves 31, 24-30, (2010).

[183] F. Bilodeau, K.O. Hill, B.Malo, D.C. Johnson, and I. Skinner, “Efficient,

narrowband LP01<—->LP02 mode converters fabricated in photosensitive

fibre : spectral response”, Elec. Lett. 27, 682, (1991).

Bibliography

133

[184] Y. J. Rao, T. Zhu, Z. L. Ran, Y. P. Wang, J. Jiang, and A.Z. Hu, “ Novel

long-period fibre gratings written by high-frequency CO2 laser pulses and

applications in optical fibre communication”, Opt. Commun., 229 , 209-

221, (2004).

[185] C-L. Lee, R-K. Lee, Y-M Kao, “Synthesis of long –period fibre gratings

with a lagrange multiplier optimization method”, Opt. Commun. 281, 61-

74, (2008).

[186] A.I.Kalachev, D.n. Nikogosyan and G.Brambilla, “Long-period fibre

grating fabrication by high-intensity femtosecond pulses at 211nm”, J.

Lightwave Tech. 23 , 2568-2578, (2005).

[187] K-W Chung , S. Yin, “Analysis of a widely tunable long-period grating by

use of an ultrathin cladding layer and higher-order cladding mode

coupling”, Opt. Lett. 29, 812-814, (2004).

[188] M. Fuumaki, Y. Phki, J. L. Brebner, and S. Roorda “ Fabrication of long-

period fibre gratings by use of ion implantation”, Opt. Lett. , 25, 88-89,

(2000).

[189] N. H. Ky, H.G. Limberger and R. P. Salathe “Efficient broadband intracore

grating LP01-LP02 mode converters for chromatic-dispersion

compensation”, Opt. Lett. 23 , 445-447, (1998).

[190] T. Nielsen , M. Fricke, D. Hellweg and P. Andresen “ High efficiency beam

splitter for multifocal multi-photon microscopy” J. Microsc. 201, 368-376,

(2001).

Appendix

134

Appendix

Dispersion relation for cladding modes

The dispersion relation equations for the cladding modes are given as [120]

)],()()()([

*)]()()()([

)]()()()()([

*)](1

)()()()([

2

1

21

2

2

32

2

1

2

2

2

321

2

32

2

2

2

2

1

21

2

2

2

1

32

2

2

2

1

21

2

2

2

32

21

2

2

2

1

2

2

222

1

2

3

2

21

2

2

322121

2

3

2

2

2

2

2

222

21

2

2

322121

2

ara

uaq

a

uapK

an

nuJ

a

uu

aran

uaq

an

uapK

an

uJ

an

uu

asun

naJraKq

n

nap

aan

uuJK

n

nu

asu

aJraKqapaan

uuJKu

vvv

vvv

vvvv

vvvv

(A.1)

The following definitions are used in above equation.

,/ 01 Zivn cleff (A.2)

Appendix

135

02 Zivn cleff , (A.3)

,11

2

1

2

2

21uu

u (A.4)

2

2

2

3

31

11

uwu , (A.5)

2,1),()/2( 2222 innu cleffii (A.6)

),()/2( 2

3

222

3 nnw cleff (A.7)

,)(

)(

111

11

'

auJu

auJJ

v

v

(A.8)

,)(

)(

233

23

'

awKw

awKK

v

v

(A.9)

),()()()()( 212122 ruYauJauYruJrp vvvvv (A.10)

),()()()()( 212

'

12

'

2 ruYauJauYruJrq vvvvv (A.11)

),()()()()( 2

'

12122

'ruYauJauYruJrr vvvvv

(A.12)

).()()()()( 2

'

12

'

12

'

2

'ruYauJauYruJrs vvvvv

(A.13)

where is the azimuthal number, Z0 is the vacuum impedance , cleffnnnn ,,, 321

represent the refractive indices of the core, cladding ,the surrounding medium and

the effective refractive index of the cladding modes in the cladding region

respectively. Jn(.) and Yn(.) are the Bessel functions of the first and the second kind

respectively. The effective indices of the m-th order cladding modes are obtained

by finding the roots of the dispersion relation in above equation.

Appendix

136

Coupling coefficient (κ) between core mode LP01 and cladding modes with

azimuthal value of 1 is given by [79] for step index fibre.

)](1

)1(

)1(*|)([1

*/)1(

*21(

2

110

11

0

111112

1

02

2

1

22

1

1

2

1

2/1

20

011

auJa

bV

bVJ

bVJauJuE

n

abVu

un

bnZ

b

cl

v

cocl

v

(A.14)

where σ is the step function for the profile of grating, 0 is the dispersion relation

for step index dual clad fibre given by

,

)()()()(

)(1

)()()()(1

2

1

2

1

21

2

1

2

1

32

2

1

2

1

21

2

2

2

32

2

2

2

222

21

2

2

322121

2

2

0

aran

uaq

an

uapK

an

uJ

an

uu

asu

aJraKqapaan

uuJKu

(A.15)

Alternatively,

Coupling coefficient k12 (= - k12 ) between LP01 and LP02 modes can be written as

[95]

,)2

exp(12

q

q

qjajkk

(A.16)

where k is given by,

,)(

)(

)(

)(*

)(

))/(1())/(1(

011

010

02

021

020

012

02

2

01

22

02

22

01

2

0

uJ

uJu

uJ

uJu

uun

VuVunkk

(A.17)

where u01 and u02 is given by,

,)4(1

)21(4/1401

V

Vu

(A.18)

Appendix

137

,1

)1

arcsin()/11arcsin(

exp2

2

02

c

c

c

c

u

V

uu

uu

(A.19)

uc is the cutoff value of the normalized frequency for LP02 mode, V (= nRk02 ) is

normalized frequency for the fibre.

Power conversion between the two modes can be written as [91]

,

2/)(

]}2/)([{sin

2

2

21

2/12

2

21

2

0

k

zkk

T

(A.20)

The power in the core and the cladding in the step index fibre are given by

,)1(

)1()1(1

2

1

112

bVJ

bVJbVJaCP ll

core

(A.21)

where C is the constant made of standard integrals associated with Bessel

functions.

The power fraction propagating in the core is given by

,cladcore

core

PP

P

(A.22)

,1)(

)()(2

1

112

bVK

bVKbVKaCP ll

clad

Appendix

138

The detuning parameter δ is given by

,2

2

101

n

cl

(A.23)

The ratio of power of with nth cladding mode and initial power at LP01 mode is

given by expression, as derived in [91]

,

)(1

)(1[sin

)0(

)(

2

22

01

g

g

gn

cl

L

P

LP

(A.24)

where L is the length of the fibre grating, kg is the coupling constant.

Spectral resonance for this case is given by,

,)(

8.0

0

2

clc nnL

(A.25)

Author’s publications

139

Author’s publications

Journals

Navin Prakash Ghimire 1, Hongchun Bao

1 and Min Gu, “Design of zero dispersive

double-clad fibre for high efficiency operation of two color light”, App. Phys. B

108:295–299 ,( 2012)

Navin Prakash Ghimire 1, Hongchun Bao

1 and Min Gu

“Broadband excitation and

collection in fibre-optic nonlinear endomicroscopy”, App. Phy. Let. 103, 073703

(2013)

Conferences

Navin Prakash Ghimire, Hongchun Bao and Min Gu, Nonlinear optical endoscopy

enabled by fibre-based dispersion compensation, 19 Australian Institute of physics

conference, Melbourne, Australia. 5-9 December 2010.

Navin Prakash Ghimire, Hongchun Bao and Min Gu, Nonlinear endoscopy using a

single zero dispersive hollow-core double-clad fibre, Focus on Microscopy-2011,

Konstanz, Germany, 17-20 April 2011.

Hongchun Bao, Navin Prakash Ghimire and Min Gu, Nonlinear optical endoscopy

for in vivo 3D super-resolution imaging, Focus on Microscopy-2011, Konstanz,

Germany, 17-20 April 2011.

Navin Prakash Ghimire, Hongchun Bao and Min Gu, Single fibre based nonlinear

endoscopy, Koala Conference-2011, Melbourne, Australia, 3-5 November 2011.

investigation of structured fibres for nonlinear ... · declaration ii declaration i, navin prakash...

Documents