Interferometric Photomechanical Spectroscopy and
Imaging of Biological and Turbid Media
by
Barry P. Payne
B.S., University of Denver (1994)
S.M., Massachusetts Institute of Technology (1997)
Submitted to the Department of Mechanical Engineering
in Partial fulfillment of the requirements for the degree of
Doctor of Philosophy
at the
MASSACHUSETTS INSTITUTE OF TECHNOLOGY
February 2001
DBarry Payne. All rights reserved.
BARKER
MASSACHUSETTS|NfN
The author hereby grants to MIT permission to reproduce
and to distribute publicly paper and electronic copies
of Ais thesis document in whole or part.
OF TECHNOLOGY
JUL 1
?001
LIBRARIES
Autho-
_________________
V
IDepartment
A
of Mechanical Engineering
December 5, 2000
Certified by
B.B. Mikic
Professor, Department of Mechanical Engineering
Thesis Supervisor
Certified by
N.S. Nishioka
Associate Professor of Medicine, Harvard Medical School
Thesis Supervisor
Accepted by
A.A. Sonin
Chairman, Departmental Committee on Graduate Students
Interferometric Photomechanical Spectroscopy and
Imaging of Biological and Turbid Media
by
Barry Payne
Submitted to the Department of Mechanical Engineering on December 5, 2000 in Partial
fulfillment of the requirements for the degree of Doctor of Philosophy
Abstract
The medical field is currently experiencing rapid growth in the area of optical
diagnostics. Minimally-invasive spectroscopic and imaging modalities enable physicians to make
increasingly accurate diagnoses in real time, without the cost and delay associated with
traditional reliance on histopathology. We have developed an ultra-high resolution
interferometric system which is well suited for clinical diagnostic applications. The
interferometric system has a spatial resolution of 0.1 nm and a temporal resolution of 3 ns.
We have utilized this high resolution interferometric system in two novel minimally
invasive techniques. Both techniques measure surface deformation of a target after absorption of
a short laser pulse. The time dependent surface deformation is a function of the target's spatially
resolved optical, thermal and mechanical properties. Therefore, accurate measurement of surface
displacement can be used to extract significant diagnostic information.
The first technique is termed Interferometric Photomechanical Spectroscopy (IPMS), and
is used to measure effective optical absorption depth of a sample. We used IPMS to measure
effective absorption depth of both diffuse and speculary reflecting targets including well
characterized colored glass samples and gelatin based tissue phantoms. The second technique is
termed Interferometric Photomechanical Tomography (IPMT), and is used to image sub-surface
absorbers such as tumors or blood vessels. IPMT combines high optical contrast with low
attenuation of sound propagation to localize sub-surface absorbers in highly scattering media.
We have used IPMT to image sub-surface blood vessels in a phantom model and in vivo.
Interferometric techniques compare favorably against diagnostic techniques which
measure surface stress instead of surface displacement. This is primarily because interferometry
is a non-contact epitaxial method capable of high resolution point measurements. Further, an
interferometric system is easily implemented into an optical fiber setup for use in minimallyinvasive catheter based procedures.
Thesis Committee
B.B. Mikic, Professor of Mechanical Engineering, MIT
N.S. Nishioka, Associate Professor of Medicine, Harvard Medical School
R. Abeyaratne, Professor and Associate Head of Mechanical Engineering, MIT
D. Rowell, Professor of Mechanical Engineering, MIT
3
-
4
1994
Acknowledgments
Now is the time to thank all the people who made life and research not only
possible, but fun and exciting. Dr Nishioka, Dr. Mikic and Dr. Venugopalan were
responsible for starting my graduate career, and for this I am extremely grateful. They
have all been great mentors and friends over the last six years, and played unique roles in
my development as a researcher.
Dr. Nishioka has always given me unlimited freedom and support during all my
different research endeavors. His easy going nature combined with his excellent practical
and clinical understanding makes him an ideal adviser for biomedical research.
Dr. Mikic has taught me so many different things regarding research and life. His
insights are always amazing, and his support has never wavered. I owe much of my
success at MIT to him. Thanks also to Liba for many wonderful evenings.
Dr. Venugopalan has always been available for detailed discussions on all aspects
of my research. His broad knowledge of the field and his ability to question key issues
has forced me pursue a deeper understanding of my own research. He was also
responsible for providing me with my first lab notebook (see opposite page).
Andrew Yablon developed the first interferometric system at Wellman
Laboratories and has always been readily available for long discussions regarding my
research ideas. His active interest and very useful insights have helped guide my
research.
I was fortunate to have both Prof. Abeyaratne and Prof. Rowell on my thesis
committee. I am very grateful for their time, and pertinent suggestions regarding my
research.
Brett Bouma and Gary Tearney have been invaluable regarding all aspects of my
interferometric setup. Their deep understanding of optics and instrumentation has not
only benefited me, but also Wellman Laboratories as a whole. In addition, they were
always willing to meet and discuss my ideas, even though they were never officially
involved in my research.
Hans Ludemann, the jua kali inspector has been a friend and colleague throughout
my PhD work. He has helped me with many aspects of experimentation and was always
willing to listen and discuss my ideas, as well as keep me thoroughly entertained. The in
vivo work was done on his forearm.
5
Stephan Brand, John Poneros and William Puricelli have been good friends
throughout my time at Wellman Laboratories. They have provided me with constant
support and entertainment.
Everyone at the Wellman Laboratories has contributed to my time here in some
way. Thank you all very much. In particular, Kevin Schomacker, Tom Deutsch, Rob
Webb, and Apostolos Doukas were always willing to discuss various aspects of my
research; Robert Redmond for the use of the laser system; Bill Farinelli for equipment
help and Susan Weeks for all her help in organizing my research assistantship.
Finally, a very special thanks to my wife and family. My family got me where I
am today, and my wife is taking me to where I am going. I owe my wife and family an
infinite amount of thanks for everything they have done for me - Asante Sana.
My wife Kelly is an amazing person, who has supported me throughout my doctoral
research. She is by far the best thing I earned as a graduate student. I am so thankful for
all the sacrifices she has made on my behalf, and I look forward to all the many
adventures in our future. This thesis is dedicated to her.
6
Table of Contents
A bstract ..........................................................................................................................
3
Acknowledgments ....................................................................................................
5
List of Figures.................................................................................................................
List of Tables ................................................................................................................
9
13
Nom enclature................................................................................................................
15
Chapter 1 Introduction............................................................................................
17
1.1 Background.............................................................................................................
1.2 Detection Type........................................................................................................
1.2.1 Detection of Input light..................................................................................
1.2.2 Detection of Em itted Light ..........................................................................
17
18
18
20
1.2.3 Detection of Heat..........................................................................................
21
1.2.4 Detection of Sound ......................................................................................
1.3 Spectroscopy and Im aging....................................................................................
21
22
1.4 Interferometry .....................................................................................................
1.5 Research Objectives ............................................................................................
23
24
Chapter 2 Interferometric Methods for Diagnostic Applications.........................27
2.1 Introduction.............................................................................................................
27
2.2 Interferometer Design ..........................................................................................
2.3 Interferom eter Capabilities ......................................................................................
2.4 Future Directions in Instrumentation ....................................................................
29
33
36
Chapter 3 Interferometric Photomechanical Spectroscopy (IPMS)......................
39
3.1 Introduction.............................................................................................................
3.2 Theoretical M odeling ..........................................................................................
3.3 Verification of Interferometric Photomechanical Spectroscopy ............................
3.3.1 Overview .........................................................................................................
3.3.2 Experimental Setup......................................................................................
3 .3 .3 R e sults .............................................................................................................
3.4 IPM S on Tissue Phantom s....................................................................................
39
42
50
50
51
55
61
3.4.1 Overview .....................................................................................................
61
3.4.2 Experimental Setup......................................................................................
62
7
3.5 Discussion...............................................................................................................
68
Chapter 4 Interferometric Photomechanical Tomography (IPTS)........................71
4.1 Introduction.............................................................................................................
4.2 Verification of Interferometric Photomechanical Tomography .............................
4.2.1 Overview .........................................................................................................
4.2.2 Experim ental Setup ......................................................................................
4.2.3 Results.............................................................................................................
4.3 Discussion.............................................................................................................
71
75
75
77
81
100
Chapter 5 Conclusions and Future Work................................................................
103
5.1 Achievem ents of This Thesis.................................................................................
5.1.1 Instrum entation..............................................................................................
5.1.2 Development and Verification of Interferometric Photomechanical
Spectroscopy..................................................................................................
5.1.3 Development and Verification of Interferometric Photomechanical
Tom ography ..................................................................................................
5.2 Future Research Directions ...................................................................................
5.2.1 Com prehensive Comparison with Pressure Transducers.................................
5.2.2 Clinical Instrum entation Developm ent ...........................................................
5.2.3 Im age Reconstruction Developm ent...............................................................
5.2.4 Sim ple Clinical Applications .........................................................................
103
104
105
106
106
106
107
107
References .................................................................................................................
109
8
104
List of Figures
Figure 2-1:
Dual balanced Interferometric system used to measure surface
displacem ent .....................................................................................
30
Figure 2-2:
A comparison of base line traces of our current interferometric
system and a previous version by Yablon and coworkers [108]........... 34
Figure 2-3:
A comparison of surface expansion traces of colored glass during
Q-switched Nd:YAG laser irradiation at 355 nm.................................
35
Simulated analytical and FEM surface displacement traces for
typical IPM S param eters. ...................................................................
50
Frequency doubled Q-switched Nd:YAG temporal laser profile with
a FW HM pulse duration of about 7 ns....................................................
52
Q-switched Nd:YAG spatial laser profile at 532 nm with a FWHM
diam eter of 6 mm ...................................................................................
52
Figure 3-4:
Schematic of experimental setup for IPMS verification.......................
53
Figure 3-5:
Absorption depth of Schott neutral density glass filters ..........................
54
Figure 3-6:
Sample displacement trace for Schott glass type NG1. ......................
56
Figure 3-7:
Sample displacement trace and best fit solution for Schott glass type
. . 57
N G l ..................................................................................................
Figure 3-8:
Sample displacement traces and best fit solutions for Schott glass
Figure 3-1:
Figure 3-2:
Figure 3-3:
type NG1. Laser fluence ranged from 12 - 440 mJ/cm 2 . ............
Figure 3-9:
Figure 3-10:
Figure 3-11:
..... ........
58
Sample displacement traces with best fit solutions for all the Schott
g lass filters.............................................................................................
59
Calculated absorption depth for 350 traces obtained under various
laser operating and alignment conditions. The axis limits
correspond to the absorption depth range given by Schott Glass
Technologies......................................................................................
61
Sample displacement trace and best fit solution for a 0.0169M
potassium chromate solution from batch A..........................................
64
9
Figure 3-12:
Figure 3-13:
Figure 4-1:
Sample displacement trace and best fit solution for a gelatin doped
tissue phantom. The tissue phantom was prepared with the 0.0169M
potassium chromate solution from batch A.........................................
66
Calculated absorption depth for 0.0169M potassium chromate
solution (batch A) and calculated absorption depth for a gelatin
tissue phantom prepared with the same solution..................................
67
IPMT scanning setup. Surface displacement traces across the
sample were obtained by scanning the tissue phantom on a digital
m icrom eter.............................................................................................
76
Figure 4-2:
Spatial profile of a focused Q-switched Nd:YAG laser beam at
1064 nm. The FWHM diameter of the spot was about 0.6 mm....... 78
Figure 4-3:
Schematic of experimental setup for IPMT ........................................
79
Figure 4-4:
Sample surface displacement of a vessel phantom model consisting
of a 1000 pm polyimide tube surrounded by pure water. ....................
83
Figure 4-5:
Sample surface displacement traces of a 1000 prm phantom vessel at
different depths. The surrounding medium was pure water................. 84
Figure 4-6:
Sample tomographic image of a 495 pm diameter sub-surface
phantom vessel in pure water. Vessel depth was 1.53 mm................... 85
Figure 4-7:
Sample contour image of a 495 pm diameter sub-surface phantom
vessel in pure water. Vessel depth was 1.53 mm. ...................................
86
Figure 4-8:
Sample contour images of three different sized phantom vessels in
pure water, The depths of the phantom vessel are all about 1.5 mm........ 87
Figure 4-9:
Normalized maximum displacement profiles for three different
phantom vessels in water. The diameters were of the vessels were
1000 pm, 495 pm and 198 pm. Gaussian functions have been fitted
to the profiles....................................................................................
. 88
Sample tomographic images for a 1000 prm phantom vessel in pure
water, 1% and 7% Intralipid solution. The depth in each media was
ab o u t 1.5 mm .........................................................................................
89
Normalized maximum displacement profiles for a 1000 prm
diameter vessel in pure water, 1% and 7% Intralipid solution.
Gaussian functions have been fitted to the profiles. .............................
90
Sample surface displacement of a vessel phantom model consisting
of a 300 prm sub-surface polyimide tube surrounded by 7 %
Intralipid solution. The vessel depth was 1.04 mm. .............................
91
Figure 4-10:
Figure 4-11:
Figure 4-12:
10
Figure 4-13:
Figure 4-14:
Sample surface displacement traces of vessel phantoms at different
depths in a 7% Intralipid solution. The phantom vessel size was 300
pm in diam eter.................................................................................
92
Sample tomographic image of a 1000 pm diameter sub-surface
phantom vessel in a 7% Intralipid solution. Vessel depth was 1.36
mm . .......................................................................................................
93
Figure 4-15:
Sample contour image of a 1000 pm diameter sub-surface phantom
vessel in a 7% Intralipid solution. Vessel depth was 1.36 mm. ............... 94
Figure 4-16:
Sample contour images of three different sized phantom vessels in a
7% Intralipid solution. The depths of the phantom vessel are all
ab ou t 1.4 m m .........................................................................................
95
Normalized maximum displacement profiles for three different
phantom vessels in a 7% Intralipid solution. The diameters were of
the vessels were 1000 pm, 495 pm and 198 pm. Gaussian functions
have been fitted to the profiles. .........................................................
96
Sample surface displacement traces of three different vessel at a
depth of 2.2 mm in a 7% Intralipid solution. The diameters of the
vessels were 1000 pm, 300 pm and 198 pm. ......................................
97
Figure 4-17:
Figure 4-18:
Figure 4-19:
Two sample surface displacement traces obtained from the forearm
of a human volunteer. The top trace was obtained directly above the
vessel, and the bottom trace was obtained to the side of the vessel.
The vessel depth is estimated to be 0.9 mm below the skin surface. ....... 99
Figure 4-20:
Sample OCT image taken directly on top of a vessel in the forearm
of a human volunteer. There is no clear indication that there is a
sub-surface blood vessel. .....................................................................
11
100
List of Tables
Table 3-1:
Material properties and calculated speed of sound for several Schott
neutral density filters..........................................................................
55
Table 3-2:
Measured absorption depths for Schott neutral density filters
compared with values obtained from Schott Glass Technologies...... 60
Table 3-3:
Measured absorption depth for three potassium chromate solutions
from two different batches (A and B) compared with estimated
. . 64
v alues. ............................................................................................
13
Nomenclature
bi
-
body force vector, [N/m 2 ]
C,
-
longitudinal speed of sound, [m/s]
CV
-
heat capacity at constant volume, [J/kgK]
D
-
characteristic penetration depth of radiation, [m]
E
-
Young's modulus, [N/m 2]
I
-
intensity of detected light, [W]
I,
-
intensity of interferometric sample arm, [W]
12
-
intensity of interferometric reference arm, [W]
k
-
wave propagation vector [radians/m]
K
-
bulk modulus, [N/M 2 ]
p
-
Laplace variable, [-]
S(t)
-
Surface displacement, [m]
so
-
Equilibrium surface displacement, [m]
t
-
time, [s]
t,
-
laser pulse duration, [s]
ui
-
displacement vector, [m]
-
acceleration vector, [m/s 2]
z
-
depth coordinate in irradiated target, [m]
15
Greek
thermal diffusivity, [m 2/s]
a
-
#
-
coefficient of thermal expansion, [K-']
3
-
Kronecker delta, [-]
Ap
-
path length difference of the interferometric system, [m]
Aw
-
AOM modulation frequency, [Hz]
el
-
strain tensor, [-]
77i
-
Lagrangian strain tensor,
01aM
-
phase of AOM signal, [radians]
0,1
-
phase of interference signal, [radians]
-
2
incident laser fluence, [J/m ]
A
-
wavelength, [m]
ya
-
radiation absorption coefficient, [1/m]
v
-
Poisson's ratio,
RI
-
3.14159...
p
-
density, [kg/m 3 ]
0-,
-
stress tensor, [N/m 2 ]
0
-
temperature relative to surrounding temperature, [K]
00
-
surface temperature relative to surrounding temperature, [K]
T,1
-
[-]
[-]
surface displacement time constant, [s']
16
Chapter 1
Introduction
1.1 Background
The medical field is currently experiencing rapid growth in the area of minimallyinvasive optical diagnostic techniques. This growth is primarily due to technological
advances in digital signal processing and optical technologies. Non-invasive
spectroscopic and imaging modalities allow physicians to make increasingly accurate
diagnoses in real time, without the cost and delay associated with traditional reliance on
histopathology. Optical techniques are attractive because they are generally safe and are
capable of extracting information on the morphological, biochemical and molecular level.
There are many different optical diagnostic techniques, each with its own advantages,
limitations, targets and resolution. We have listed several recent review articles that cover
a broad range of current optical diagnostic techniques [31, 78, 81, 93, 99].
All optical diagnostic techniques use light with wavelengths ranging from the
ultraviolet to the infrared to obtain tissue information. These techniques differ in how this
17
information is extracted, and the type of information extracted. How the information is
extracted is a function of what is specifically measured. There are four distinct types of
detection. These are the detection of input light, detection of emitted light, detection of
heat and detection of sound. These four detection types are reviewed in section 1.2. The
type of information extracted is either spectroscopic based, imaged based, or a
combination of both. Spectroscopy is primarily concerned with wavelength dependent
properties of tissue. Imaging is concerned with spatially mapping tissue structure or
function. Section 1.3 reviews two important sub-sections of spectroscopy and imaging;
tissue optical properties, and imaging in highly scattering media. Understanding both
how information is extracted, and what information is extracted is key in developing new
optical diagnostic applications and improving existing ones.
1.2 Detection Type
1.2.1 Detection of Input light
The detection of input light can be further subdivided into the detection of
coherent (ballistic) light, and the detection of diffuse light. The detection of coherent light
is normally associated with diffraction limited imaging applications. The main challenge
for techniques based on the detection of coherent light is to reject non-coherent or
scattered light. The most common methods to reject scattered light are spatial filtering,
polarization gating and time gating.
Spatial filtering is one of the simplest and oldest methods for suppressing
scattered light. The concept is to use a spatial filter to block scattered light, while
allowing the majority of coherent light to be detected. The most common example of a
18
spatial filter is a simple pin hole which is most often used to image point objects. The
confocal microscope is the most widely used instrument that uses spatial filtering for high
resolution imaging (e.g. [72]).
Polarization gating utilizes polarization differences between scattered light and
unscatterd light to reject the former. Coherent light retains its polarization state, whereas
scattered light is randomly polarized. Polarization gating is generally performed by
resolving the polarization state of light emerging from a sample, and using this
information to subtract the diffuse background (e.g. [43]).
Time gating uses the flight time of photons to obtain path length information. In
the simplest configuration, a time resolved measurement of light emerging from a sample
that was illuminated by a short input pulse provides a temporal distribution of path
lengths. Coherent light has the shortest flight time and is temporally separated from
multiply scattered light. Simple time gating can be achieved by a streak camera [22, 37].
Improved time gating can be achieved with correlation techniques where the sample light
is mixed with a short reference pulse of proper time delay. Nonlinear correlation
techniques include SHG [25, 56, 110, 111], upconversion [27], Kerr gating [104], Raman
amplification [6] and Photorefractive holography [47]. Linear correlation techniques
include time-gated holography [61] and optical coherence tomography [13, 31, 83].
Optical coherence tomography is the most clinically successful linear correlation
technique and combines heterodyne detection and low coherence interferometry. OCT
was first demonstrated in the eye [44] and has since been used in many clinical
applications [7, 10, 12, 14, 29, 30, 49, 50, 54, 75, 84, 85, 88, 91, 96].
19
The detection of diffuse light is normally used to measure tissue optical properties
or to perform low resolution imaging. The optical properties or internal structure of tissue
is reconstructed from diffuse transmittance or reflectance measurements by using a
theoretical model of wave propagation. In general, the diffusion approximation to the
radiative transport equation provides sufficient accuracy, but needs to be supplemented
with complex boundary conditions [5]. Diffuse light detection is performed in either the
spatial, time or frequency domain. In the spatial domain or spatially resolved method,
light of constant intensity is launched into tissue and detected at different known spatial
locations [20, 28, 45]. In the time domain or time resolved method, a pulse of light is
launched into tissue. Changes in the pulse shape due to attenuation and scattering are
used to extract optical absorption and scattering properties. One source-detector
separation is sufficient to characterize the optical properties [51, 64, 70]. The third
approach uses high frequency (MHz) intensity modulated light to create diffuse photon
density waves (PDW). The PDW propagate with a wave vector that is a function of
optical properties [15, 32, 71, 74, 86, 100].
1.2.2 Detection of Emitted Light
The detection of emitted light covers all fluorescence based techniques. The
principles of fluorescence spectroscopy and fluorescence imaging are well understood
(e.g. [59, 105]). The advantage of fluorescence based techniques is the potential to link
biochemical and morphological properties of tissue to individual patient care. There are
numerous fluorescence techniques and configurations which are based on one, two and
three photon excitation, confocal microscopy, fluorescence life times as well as a wide
20
range of both endogenous and exogenous fluorophores [24, 39, 40, 72, 78, 79, 81, 90, 92,
103, 107].
1.2.3 Detection of Heat
The detection of heat (temperature) during and after laser irradiation can be used
to extract tissue optical properties and to perform low resolution imaging. For example,
pulsed photothermal radiometry measures the temporal evolution of radiative emission
from the surface of a tissue sample after the absorption of a sub-ablative laser pulse. This
measurement is used to compute the original thermal field, thereby extracting optical or
structural information [53, 62, 63, 65, 66, 102]. This is an inverse diffusion problem and
it is therefore very difficult to accurately reconstruct the original thermal profile.
1.2.4 Detection of Sound
The detection of sound that is generated from the absorption of light is a
potentially powerful technique for extracting the optical properties of tissue, or imaging
sub-surface structures like blood vessels or tumors within tissue. The generation of sound
by light absorption is often termed the photoacoustic or optoacoustic effect. Detailed
theoretical and experimental aspects of photoacoustic processes are listed in three
excellent references, including a review paper, book chapter and book [36, 46, 87].
Although the field of photoacoustics is well established, there has been a recent trend
towards applying photoacoustic techniques for medical diagnostic applications. These
techniques include photoacoustic spectroscopy and photoacoustic imaging [26, 41, 55,
58, 67, 68]. Although photoacoustic spectroscopy and imaging are inverse problems,
21
where the original stress distribution is reconstructed from time dependent stress
measurements, it is not prone to the same instabilities and inaccuracies of diffusion based
inverse problems such as pulsed photothermal radiometry. This is simply because wave
propagation is not significantly affected by turbid media, as is the case for diffusion
based light propagation. This is a major advantage for imaging in highly scattering media
such as tissue.
1.3 Spectroscopy and Imaging
The optical properties of tissue is an important sub-section of spectroscopy, and
play a central role in many diagnostic applications. Tissue optical properties are
responsible for the propagation of light in tissue, and can also yield useful diagnostic
information regarding metabolic, physiologic and structural status of the tissue [18, 106].
For example, changes in optical properties have been used to study the hemodynamics of
brain, [35, 89, 98]. The fundamental optical properties of tissue are the absorption
coefficient, scattering coefficient, and scattering phase function or scattering anisotropy.
Often the scattering coefficient and phase function are combined to form a reduced
scattering coefficient. For one dimensional geometry, the penetration of light decreases
exponentially, and an effective penetration depth is used which is a function of the
absorption and reduced scattering coefficient.
Imaging modalities capable of imaging deep sub-surface absorbers in highly
scattering media would be very useful in many diagnostic applications such as breast
cancer screening. Coherent imaging techniques such as optical coherence tomography
and confocal microscopy cannot be used to detect deep absorbers since tissue scattering
22
is too great. Currently, non-optical imaging techniques such as X-ray radiography,
magnetic resonance imaging and ultrasound imaging are used to image deep sub-surface
absorbers. Although these techniques are valuable, the major limitation is usually
associated with low contrast, leading to insufficient sensitivity for detection of small
tumors.
1.4 Interferometry
Interferometric methods have shown great promise in a variety of diagnostic
applications. The success of interferometric techniques is primarily due to the high spatial
and temporal resolution that can be achieved. In addition, an interferometric system is
easily implemented into an optical fiber setup for use in minimally-invasive catheter
based procedures. For example, optical coherence tomography has been successfully
during endoscopic procedures. (e.g. [12, 85, 95] ).
Interferometry has also been successfully used in pump-probe techniques where
the thermo-mechanical response of biological tissue was investigated by measuring the
time dependent surface displacement of biological tissue after the absorption of a short
laser pulse [4, 21, 73, 82, 108]. The time dependent surface displacement is a function of
the sample's optical, thermal and mechanical properties. Interferometric monitoring of
surface deformation after laser irradiation is therefore a potentially powerful diagnostic
tool. Stress transducers could also be employed as a diagnostic tool to measure material
properties [55, 58, 67, 68]. However, stress transducers are not ideal in many medical
applications since they are contact based methods and are not easily used in conjunction
with minimally invasive catheter type procedures.
23
1.5 Research Objectives
Given the importance of tissue optical properties, and the need for sub-surfacing
imaging in scattering media, the goal of this thesis was to develop a minimally-invasive
spectroscopic technique, and a minimally-invasive imaging technique. Interferometry is a
non-contact minimally invasive high resolution technique and is well suited for both
these diagnostic applications. The specific objectives of this work are listed below:
I.
Develop a high resolution interferometric system capable of accurately measuring
surface displacement on both diffuse and specular targets
II.
Develop an accurate spectroscopic technique to measure the effective absorption
depth of well characterized targets and tissue phantoms
III.
Develop an interferometric technique to image sub-surface absorbers such as
blood vessels in highly scattering media.
Given these objectives, the rest of this thesis is organized into four sections. Chapter 2
describes the interferometric system developed and illustrates its capabilities as compared
to previous interferometric systems. Chapter 3 describes a novel minimally invasive
spectroscopic technique termed Interferometric Photomechanical Spectroscopy (IPMS).
IPMS was used to accurately measure the optical properties of colored glass as well as
tissue phantoms. Chapter 4 describes a novel minimally invasive imaging technique
24
termed Interferometric Photomechanical Tomography (IPMT). IPMT was used to image
sub-surface blood vessels in a phantom model and in vivo. Finally, Chapter 5 summarizes
this thesis and discuses future directions of this work.
25
Chapter 2
Interferometric Methods for
Diagnostic Applications
2.1 Introduction
Interferometric methods are attractive for non-invasive diagnostic imaging
applications. In particular, heterodyne interferometry, a linear correlation technique
capable of shot noise limited detection has shown great promise in a variety of medical
imaging applications [81]. The success of interferometric techniques is primarily due to
the high spatial and temporal resolution, and the ability to be implemented within an
optical fiber system. High resolution imaging through optical fibers is valuable for
minimally invasive catheter based diagnostic applications. For example, Optical
Coherence Tomography (OCT) is a very promising minimally invasive interferometric
technique used for high resolution imaging of biological tissue [31, 83]. The basic
principle of OCT is analogous to ultrasound, where echoes from sub-surface structures
are used to create tomographic images. OCT uses echoes of light rather than echoes of
sound, and therefore has superior resolution to ultrasound. OCT was first demonstrated in
27
vitro in the human retina and atherosclerotic plaques in 1991 [44]. The first human
studies with OCT were performed in 1993 in ophthalmology [29, 91], and have expanded
to include dermatology [84], dentistry [7], cardiology [14, 30], urology [54, 96],
gynecology [10] and gastroenterology [12, 49, 50, 75, 85, 88].
Interferometry has also been successfully used in time resolved pump-probe
techniques to study the thermo-mechanical response of biological tissue [4, 21, 73, 82,
108]. These techniques measured the time dependent surface displacement of biological
tissue after the absorption of a short laser pulse. The time dependent surface displacement
is a function of the sample's thermo-mechanical properties as well as the specific lasertissue geometry such as spot size. The thermo-mechanical response of biological tissue
can be used to extract useful diagnostic information about the tissue's optical, thermal
and mechanical properties. Photoacoustics or optoacoustics is an analogous technique
which uses stress instead of displacement to extract the optical properties of tissue [55,
58, 67, 68]. Stress detection via a pressure transducer is not ideal in many medical
applications since it is a contact method and is not easily used in conjunction with
minimally invasive catheter based procedures.
The success of interferometric systems in medical applications is highly
dependent on the temporal and spatial resolution of the instrument. Previous
interferometric systems employed to measure surface displacement had temporal
resolutions of 3-4 ns, and spatial resolutions of 1-4 nm [4, 21, 73, 82, 108]. Although the
temporal resolution is adequate for most applications, the spatial resolution is not
sufficient for diagnostic applications which must differentiate subtle changes in tissue
properties, or imaging applications attempting to image deep sub-surface absorbers.
28
We have developed an ultra-high resolution heterodyne interferometric system
capable of angstrom spatial resolution, and sub nanosecond temporal resolution. The high
spatial and temporal resolution is ideal for measuring surface displacements with
extremely high accuracy. In Chapter 3, we have used this interferometric system to
measure optical absorption depth of well characterized glass samples and tissue
phantoms. In Chapter 4, we have used this interferometric system to image sub-surface
absorbers within tissue phantoms and in vivo.
2.2 Interferometer Design
The interferometric system developed and employed in this thesis is shown in
figure 2-1. This configuration is based on a previous version by Yablon and coworkers
[108], with two important differences. First, the focusing lens was replaced with a
gradium index lens with a focal length of 80 mm. The gradium index lens maintained
better spatial coherence in the sample beam, and therefore enabled a higher interference
depth of modulation and lower noise. The interference depth of modulation is the ratio of
interfered light to the total available light. A depth of modulation of 100% means that all
the available light is being used for interference, implying perfect spatial and temporal
coherence. The shorter focal length of the gradium index lens reduced the probe spot
diameter to about 10 pm, which enabled more coherent light to be reflected back off
diffuse surfaces. Although a smaller focal length lens could be used to decrease the spot
size further, the sample working distance would be reduced which is not practical for
most applications.
29
Second, the avalanche photodiode was replaced with dual balanced Si PIN
photodiodes. The Si PIN photodiodes have very high temporal resolution (0.2 ns) and
very low noise owing to the small dark current (2 pA). The dual balanced detection
system reduces the effect of local oscillator noise by subtracting their photocurrents to
cancel unwanted DC components [1]. Ideally, a transimpedance amplifier should be used
to amplify the subtracted photocurrents, although a broadband low noise amplifier is
sufficient.
X/4 plate
Dielectric
Mirror
Incident laser
Pump beam
Polarized
Beamsplitter
Unshifted
Beam
Ea
Dielectric
Beam
Expande r
Mirror
Polarized
HeNe Laser
A
=632.8 nm
Frequency
Shifted
Ba
Lplate
X/2
Focu sing
Lens
AcousticOptical
modulator
Non-Polarized
Beamnsplitter
Dielectric
Mirror
Target
Dual Si PIN
Detectors
Figure 2-1:Dual balanced Interferometric system used to measure surface displacement.
30
A polarized 3mw helium-neon (HeNe) laser beam (Melles Griot, Irvine, CA) with
a wavelength of 632.8 nm is used as the interferometric probe beam. An acoustic optical
modulator (IntraAction Corp., Bellwood, IL) is used to deflect and impart a frequency
shift of 110 MHz to the reference beam. The unshifted beam, also called the sample
beam, passes through a polarized beamsplitter and quarter wave plate and is directed
towards the target by a dielectric mirror. The quarter wave plate rotates the polarization
state by 45'. When necessary, a
loX
beam expander (Melles Griot) and 80 mm focal
length gradium index lens is used to increase the numerical aperture, yielding a spot size
of about 10 pm. This is useful for imaging diffuse surfaces such as gelatin phantoms and
tissue targets. The HeNe laser reflects off the sample surface due to a refractive index
mismatch between the air and target, and passes back through the quarter wave plate,
rotating the polarization state by an additional 45'. The polarization state of the sample
beam is therefore rotated by a total of 90', causing the beam to be reflected by the
polarized beamsplitter, rather than transmitted. The reflected sample beam passes through
a half wave plate to restore the original state of polarization and is recombined with the
reference beam by a 50/50 non-polarizing beamsplitter. The recombined beams have a
900 phase difference and are imaged by 25.4 mm focal length lenses on two Si PIN
photodiodes (Hamamatsu, Bridgewater, NJ). The small active area of each PIN diode
(0.126 mm 2 ) enables confocal imaging of the target. The photocurrents generated by the
Si PIN photodiodes are subtracted, canceling any inherent laser noise. The subtracted
photodiode current is amplified by a commercial 500 MHz broadband low noise
amplifier (Mini-Circuits, Branson, MO) and is digitized by a 500 MHz bandwidth
oscilloscope (Tektronix, Wilsonville, OR). The output of the commercial low noise
31
broadband amplifier was compared to a lab built transimpedance amplifier. The signal to
noise ratio for the transimpedance amplifier was slightly better than the signal to noise
ratio of the broadband amplifier, but the gain curve was not as flat. A commercial
broadband transimpedance amplifier would probably yield better performance than a
commercial low noise broadband amplifier.
When the target is stationary, the photodiodes detect interference fringes that
oscillate at the AOM drive frequency of 110 MHz. When the sample surface moves, the
path length difference between the two arms changes, and the phase of the modulation
frequency is altered. The polarity of the phase shift depends on the direction the sample
moves. The intensity reaching the photodiodes is a function of time (t), path length
difference (Ap), and intensities of the sample and reference arms and is given by:
I(Ap,t)= I , + I2 + 2
cos(o,,).
(2.1)
I1 and I2 are the intensities of the sample arm and reference arm, and 0,' is the phase of the
modulated signal which is given by:
0, = kAp+ Awt.
(2.2)
k is the propagation vector (2zc/A) and Aw is the AOM modulation frequency. The surface
displacement S(t) is equal to half the path length difference and given by:
S(t) -
_Ap _-,
2
_-
-
Awt
2k
_
-
-
pAOM
.
2k
32
(2.3)
The surface displacement is therefore computed by subtracting the phase of the
AOM drive signal from the phase of the detected signal. The phase of both signals are
obtained with the aid of the Hilbert Transform. In general, the AOM signal is sufficiently
pure that subtraction of a simulated AOM phase is equivalent to subtracting the phase
from the real AOM signal.
2.3 Interferometer Capabilities
The improvements of our interferometric system over a previous version
developed by Yablon and coworkers [108] is best illustrated by comparing sample
baseline traces from both systems as shown in figure 2-2. Both baseline traces were
obtained under similar conditions and filtered with a 80MHz band pass filter resulting in
a temporal resolution of 4 ns. The spatial resolution defined by the standard deviation
(STD) of the displacement signal has been improved by an order of magnitude.
33
2
- Previous system
STD = 0.93 nm
0
.... . .
- - - - -.-
--
Crent version
-20
0
--
-
-
-
--
-
STD
100
200
300
400
=
0.09 nm
500
Time (ns)
Figure 2-2:A comparison of base line traces of our current interferometric system and a previous
version by Yablon and coworkers [108].
Two surface expansion traces of colored glass after Q-switched Nd:YAG laser
irradiation at 355 nm are shown in figure 2-3. One trace was obtained with our new
interferometric system, and one trace was obtained with the previous interferometric
system. The maximum displacement in both cases was about 30 nm, although they have
been normalized for display purposes. The significant improvement in spatial resolution
is highly desirable for many diagnostic applications. For example, subtle changes in the
optical absorption depth may allow contrast between pathological and non-pathological
structures.
34
1.2
0 .8 - - - -- - -- -0.8Current
---- --- -- --
0.2 -
-0.2-200
---- --
version
Previous system
-
-
0.4 -
--
---
-
Laser pulse on
-100
0
100
200
300
400
Time (ns)
Figure 2-3:A comparison of surface expansion traces of colored glass during Q-switched
Nd:YAG laser irradiation at 355 nm.
In general, there is a trade off between the temporal and spatial resolution of the
interferometric system. For the experiments in this thesis, a 80 MHz digital band pass
filter was used resulting in a 4 ns rise time. The best STD achieved on a highly reflecting
target without using the
loX beam expander or focusing lens was 0.08 nm. This spatial
resolution represents the limit of the 8 bit digitizing oscilloscope. For samples that reflect
only about 4% of the incident light such as glass, the STD increases slightly to 0.09 nm.
When the beam expander and focusing lens were used, the interference depth of
modulation decreased from 98% to about 75%. This meant there was some spatial
incoherence between the reference and the sample beams caused by imperfections of the
35
beam expander optics. Spatially incoherent light incident on detectors contributes to
noise. Consequently the best STD achieved on a high reflecting surface was 0.09 nm, and
the best STD achieved on a low reflecting sample such as glass was to 0.12 nm. The
small increase in STD due to spatial incoherence highlights the benefit of using low noise
detectors such as Si PIN photodiodes.
For diffuse surfaces such as gelatin tissue phantoms, the amount of coherent light
reflected off the surface decreases and the STD increases. The best STD achieved on a
gelatin tissue phantom was 0.16 nm.
The temporal resolution of the interferometric system is limited only by the 500
MHz bandwidth of the digitizing oscilloscope and amplifier. Therefore, sub-nanosecond
rise times are achievable with the current system, although the spatial resolution will
degrade. For example, the STD for a temporal resolution of 0.9 ns on a highly reflecting
surface was 0.28nm. Since the Si PIN photodetectors have a bandwidth of 1.5 GHz, one
could in theory measure rise times as low as 230 ps.
2.4 Future Directions in Instrumentation
There are two distinct tasks for improving the interferometric system described in
this chapter. The first task involves improving existing components in the current system,
and the second task involves transforming the current system into a fiber based clinically
useable tool.
The major limiting component in the current system is the HeNe probe laser
which has insufficient power to enable shot noise limited detection on low reflecting
surfaces. Ideally, a dual balance detection system should be operated just under the
36
threshold of the photodetectors. This was not possible with our current system for low
reflecting targets. For highly reflecting targets, noise calculations indicated that detection
was limited by shot noise. These calculations were based on a transimedance amplifier
configuration, but should be comparable to a broadband low noise amplifier. A higher
power HeNe laser would allow shot noise limited detection even on diffuse surfaces. The
specific laser choice must be made carefully. Generally, higher power lasers have a
longer cavity length, and therefore shorter beat frequency [38]. If the beat frequency is
too close to the carrier frequency, the interference phase is corrupted and spatial
resolution is significantly reduced. Although filter techniques will reduce this effect, we
feel shorter cavity higher power lasers are best suited.
The next major improvement would be to implement this system into a optical
fiber based system. This would not only ease alignment, increase the depth of
modulation, but also enable the current system to be used in real clinical applications,
including minimally invasive catheter based procedures. The authors feel that
implementing the current system into an optical fiber setup should not be difficult. This is
because interferometric fiber based technologies already exist for other techniques like
optical coherence tomography.
Finally, the clinical value of the interferometric system would be greatly increased
by enabling the probe beam and pump beam to be scanned. This would allow multiple
spatial displacement traces to be obtained quickly. This is particularly useful for imaging
purposes. Alternatively, one could envision a probe array for collecting multiple points
simultaneously. This would avoid scanning difficulties and allow 3-dimensional
reconstruction of sub-surfaces absorbed in a single shot.
37
Chapter 3
Interferometric Photomechanical
Spectroscopy (IPMS)
3.1 Introduction
The optical properties of tissue play a central role in many therapeutic and
diagnostic medical applications. For example, therapeutic applications such as laser
surgery and photodynamic therapy rely on predicting the spatial distribution of laser
energy within the tissue, and hence ensuring correct laser dosimetry. The spatial
distribution is a function of the absorption and scattering properties of tissue. Tissue
optical properties also yield useful diagnostic information regarding metabolic,
physiologic and structural status of the tissue state. In many applications, it is often
convenient to characterize the spatial energy distribution for a large diameter beam by an
effective absorption depth. The effective absorption is a function of absorption
coefficient, scattering coefficient and an anisotropy parameter which governs the angular
dependence of scattering [18, 106].
39
There are a number of methods used to measure the effective absorption depth of
a sample. These methods either directly measure optical properties via simple attenuation
measurements, or indirectly extract optical properties via a theoretical model. In general
indirect methods are based on the diffusion approximation to the radiative transport
equation and measure the diffuse optical transmission and reflectance of a sample at
different wavelengths. These techniques are performed in either the spatial, time or
frequency domain. In the spatial domain or spatially resolved method, light of constant
intensity is launched into tissue and detected at different known spatial locations [20, 28,
45]. In the time domain or time resolved method, a pulse of light is launched into tissue.
Changes in the pulse shape due to attenuation and scattering are used to extract optical
absorption and scattering properties. One source-detector separation is sufficient to
characterize the optical properties [51, 64, 70]. The third approach uses high frequency
(MHz) intensity modulated light to create diffuse photon density waves (PDW). The
PDW propagate with a wave vector that is a function of optical properties [15, 32, 71, 74,
86, 100].
Recently, photoacoustic or optoacoustic techniques have been used to extract
optical properties of turbid media [55, 58, 67]. Photoacoustic spectroscopy extracts the
optical properties of a sample by measuring thermoelastic stresses caused by the
absorption of a short laser pulse. Thermoelastic stresses are a function of the initial
energy distribution within the sample, and therefore also a function of the sample's
optical properties. Although photoacoustic spectroscopy is an inverse problem, where
optical properties are extracted by reconstructing the original stress distribution, it is not
prone to the instabilities and inaccuracies of diffusion based inverse problems. This is
40
simply because wave propagation though turbid media is not altered significantly
compared with diffusion based propagation. Consequently, it is possible to measure
optical properties of structures deep inside a sample, where photon diffusion based
techniques fail.
Current photoacoustic spectroscopy techniques employ pressure transducers to
measure laser induced stresses. In most cases, the measurement must be performed in an
epitaxial or so called backward mode since access to both sides of a tissue target is not
possible. Standard front surface pressure transducers cannot be used as they would
physically block the incident laser light. Investigators have therefore used a variety of
creative techniques to overcome this problem, including piezoelectric transducers with
separated light and sound fields, transparent transducers measuring optical reflectance,
annular piezoelectric elements, and acoustic conductors where sound generated by an
obliquely incident laser pulse propagates predominantly normal to the tissue surface [8, 9,
17, 55, 57, 69]. Despite these technologies, stress detection via a pressure transducer is
not ideal in many medical applications since it is a contact method and is not easily used
in conjunction with minimally invasive catheter based procedures.
In this Chapter, we have developed a novel minimally-invasive non-contact
interferometric technique used to measure the effective absorption depth of a sample with
extremely high accuracy. We call this technique Interferometric Photomechanical
Spectroscopy (IPMS). We use the word photomechanical instead of photoacoustic or
optoacoustic since we measure displacement and not stress. IPMS uses an interferometric
system to measure surface expansion during and after the absorption of a short laser
pulse. The expansion of the sample is caused by thermoelastic stress relaxation and the
41
time constant of this relaxation is proportional to absorption depth. The magnitude of
expansion is a function of incident laser energy and thermo-physical properties of the
sample. This technique is complementary to a similar technique called Interferometric
Photothermal Spectroscopy (IPTS) [108, 109]. IPTS calculates effective absorption depth
by measuring surface cooling due to thermal diffusion.
The concept of using an interferometric based system to calculate absorption
depth by monitoring surface expansion has been noted previously, although it was not
termed IPMS nor was it employed to measure an unknown absorption coefficient [2, 4,
48] [3]. Recently, the same group made some preliminary absorption depth
measurements on meniscus samples [21]. However, IPMS has yet to be fully verified, or
its accuracy accessed. In addition, the spatial resolution of the interferometric system
used in these experiments is 50 times worse than the spatial resolution of our
interferometric system.
3.2 Theoretical Modeling
In this section, we derive the one dimensional relationship between surface
displacement and absorption depth using the theory of elasticity. A non-linear fitting
algorithm is used to extract the absorption depth from the displacement traces acquired
by the interferometric system. We begin by deriving Green's solution which assumes the
incident laser energy is absorbed instantaneously. The following derivation was based on
two references [60, 97].
The stress and strain relationship in any material are governed by three equations;
the stress tensor, the equation of motion and the constitutive relation for that material. We
42
neglect effects of thermal diffusion since the time scale of stress relaxation is much
smaller than the time scale for thermal diffusion. The strain tensor (Eij) is given by eq.
(3.1) and is the small strain approximation of the Lagrangian strain tensor (7i). The
subscripts represent indicial standard notation.
77'; ~ Ej =
(3.1)
Ui + Uji,
where u1 is the displacement vector. The equation of motion is derived from the
conservation of momentum and is given by:
ai
(3.2)
+ b, = p 5,
where ai, is the stress tensor, bi is the body force vector, p is the density and Vj is the
acceleration vector. The constitutive relation for a linear isotropic material is given by:
(3.3)
+v
E
E
where v is Poisson's ratio, E is elastic modulus,
o is the relative temperature
# is the coefficient of thermal expansion,
and 3 is the Kronecker delta. Assuming the laser energy is
absorbed exponentially, the temperature distribution will be:
43
S=
ee
,D
(3.4)
where 0 is the surface relative temperature, and D is the effective absorption depth which
is equal to the inverse of the absorption coefficient. Substituting eq. (3.4) and (3.3) into
eq. (3.2), we obtain the partial differential equation for displacement as a function of
depth (z) and time (t).
-U2
02
a2
/KQ
2
2
dZ2
P,
zu2
"
Dpo
(3.5)
e
where c, is the longitudinal speed of sound and is defined in eq. (3.6) and K is the bulk
modulus and is defined in eq. (3.7).
2 -
/
E(l - v)
p(l+ v)(l - 2v)
(3.6)
.E
E
3(1 - 2v)
(3.7)
and
K=
The initial conditions and boundary conditions are as follows:
44
u=O
du
t =0,
-o
t =0,
dt
u=0
Z
(3.8)
-+4o
and
-=O
z=0.
The solution is conveniently obtained by using the method of Laplace transforms. The
Laplace transform is obtained by the following integral transform:
L{f(t)} = f(p) =
e-P'f(t)dt.
(3.9)
Transforming eq. (3.5) into the Laplace domain yields
2-
q U=
d_2_
0i
2=
+-e
(3.10)
,
P
where
Dpc
and
(3.11)
q=-P.
C1
The general solution is given by:
U = Cie- q
+ c2e ''z +
{D2
QD
2 2
p .q D2-1
45
1e- D.
(3.12)
Solving for the constants of integration c, and c2 by applying the boundary conditions in
(3.8) yields the final solution in the Laplace domain:
Us-== QDc
,2
2_eY'2
e
p9
-
+
2
QD 2,Y2
p(p 2
2e
_7
D
(3.13)
2
y )
where
CD
(3.14)
D
The final solution is transformed back into the time domain using the following
transforms:
L
t<a
0,
f(t
-
a),
t &aJ
_
I- =
P
= e- f(P),
and
The one dimensional surface displacement is finally given by:
KQD [
U = #PODe
T'- I
2 p-c2
'*i)
z
-2e
(3.16)
C1
and
KK0D
e
e
"
2=
+e
z
-2e1
t<-.
C1
2pI
46
(3.17)
p+ca
= e'
(3.15)
At the surface (z=0), the solution simplifies to:
i
S(t)=-u.
=
,
(3.18)
So 11-e
where
SO _- PKDC 2
S
p1f
,,D 1I+v
3
Ip+ V.
(1-v)
3pc, (I-
(3.19)
v.
0, is the incident laser fluence, and c, is the heat capacity at constant volume. The time
constant of the surface displacement is a function of absorption depth D and longitudinal
speed of sound c, only: r,, =D/c,.
In general, the input laser pulse has a finite duration and the impulse solution
must be convoluted with the laser pulse temporal profile. The surface displacement can
be calculated from the following equation:
S(t) = f (I - e-")g(x)dx,
(3.20)
where d = T-' and g(x) is the laser temporal profile. The temporal profile of the laser
system used in this thesis resembled a Gaussian function given by:
(3.21)
g(x) = Ie-",
47
where
2.8
.
FWHM2
a =F W
(3.22)
FWHM is the full width half max of the laser temporal profile and I is the peak intensity.
Solving eq. (3.20) yields the final one dimensional surface displacement caused by the
absorption of a Gaussian laser pulse:
IQ0 +v
S(t) = 3
-/"
pc, ( I-v
4a
+erf a2 t -e
dc42
4a
e-1dtjIl+erf1.
2at -d
4a
(3.23)
For short laser pulses or long absorption depths (t, << r,,) eq. (3.23) will simplify
to the impulse solution given in eq. (3.18). For long laser pulses or short absorption
depths (t >>)
the surface displacement will simply replicate the integral of the incident
laser pulse profile. In the case of a Gaussian profile, the surface displacement will be
given by eq. (3.24). Note, the onset of surface expansion in both eq. (3.23) and eq. (3.24)
occurs before t = 0.
S(t) =
&
+erf a2t-.
(3.24)
48
It is interesting to note that eq. (3.24) could be used to study photochemical
processes such as fluorescence lifetimes, where there is a delay thermalizing absorbed
laser energy.
For all of our experiments, the effect of the finite laser pulse was small since the
laser pulse duration t, was much shorter than the time constant -,, of the surface
displacement. This condition is termed stress confinement and is defined as t << r,.
The governing equations were also implemented numerically using a finite
element method (FEM) partial differential equation solver (Matlab, The Mathworks,
Natick, MA). The FEM solution and analytical solution were identical and are both
shown in figure 3-1. The properties used for the figure correspond to typical parameters
used in the IPMS verification experiments.
49
S 0.2 -
D =100 pm
-----
C = 5.5 pm/ns
FWHM = 8 ns
-50
0
50
100
150
Time (ns)
Figure 3-1:Simulated analytical and FEM surface displacement traces for typical IPMS
parameters.
3.3 Verification of Interferometric Photomechanical
Spectroscopy
3.3.1 Overview
To test the validity and access the accuracy of IPMS, a set of well characterized
glass samples were investigated. The glass samples chosen were absorbing neutral
density filters obtained from Schott Glass Technologies Inc. (Duryea, PA). Neutral
density filters were chosen because the absorption depth does not vary significantly with
50
wavelength. In particular, the absorption depth between 450 nm and 600 nm remains
essentially constant.
3.3.2 Experimental Setup
The interferometric system was previously described in Chapter 2 and illustrated
in figure 2-1. The beam expander and 80 mm gradium focusing lens were not necessary
since the glass targets have smooth surfaces. The larger interferometric spot size
compensated for minor spatial inhomogeneities within the incident pump laser beam and
avoided the need for complex beam homogenizing techniques.
Laser System
The pump laser used was a Q-switched Nd:YAG (yttrium aluminum garnet with
neodymium ions) laser, frequency doubled to operate at 532 nm. This wavelength was
chosen since it is centered in the flat portion of the absorption spectrum of the neutral
density filters. The pulse duration of the laser beam was about 7 ns, and had a Gaussian
temporal shape. figure 3-3 illustrates a typical temporal profile. The spatial profile was
analyzed using a CCD Camera (Cohu 4800, Cohu Inc), and found to resemble a flat top.
The full width half max (FWHM) diameter of the raw beam was estimated to be 6 mm.
Figure 3-4 illustrates the spatial profile.
51
0.9
----------------
*
........--------
*
0.8
**
0.7
-*
0)
0.6
- -
0.5
0)
-
-
-
-
-
-
-
-
- -
-
-
-
- - -
-
. .
..
. .
.
*
-
-
*
.
. .. . . . . . . . . . . .
0.4
...
. .... . .
. .
.
. .
. .*.
.
0.3
. .. .... . ..
. ..
*
.
-.
.
.
* . . . . . . . . ..
-
- - -
-
-
-
-
-
-
0.2
-
0.1
0
-3'0
-20
-10
10
0
20
r
30
Time (ns)
Figure 3-2:Frequency doubled Q-switched Nd:YAG temporal laser profile with a FWHM pulse
duration of about 7 ns.
002
6-
Distance (mm)
2
0
0
2
Distance (mm)
Figure 3-3:Q-switched Nd:YAG spatial laser profile at 532 nm with a FWHM diameter of 6 mm.
52
The frequency doubled Nd:YAG laser irradiation was directed toward the sample
by dielectric mirrors. The angle of incidence at the sample was 15'. The angle of laser
radiation propagation within the target is a function of the angle of incidence and the
refractive index of the target. The refractive index for glass is about 1.5, and so the angle
of laser radiation propagation within the glass targets was about 100. The measured
absorption depth was corrected for this angle by dividing by the cosine of 10'. A diagram
of the setup is shown in figure 3-4. The laser fluence (energy per unit area) used for
experiments ranged from 10 - 550 mJ/cm 2. It should be noted though, that IPMS does not
rely on incident laser fluence to calculate absorption depth.
Interferometric
system
' Interferometer
Dielectric
mirror
probe beam
Nd:YAG
532 nm
840(
Figure 3-4:Schematic of experimental setup for IPMS verification
53
Glass Sample
Glass Targets
The absorption depth for several neutral density filters is shown in Figure 3-5.
The absorption depth was calculated from transmission curves supplied by the
manufacturer, and under the assumption that the laser energy is absorbed exponentially
with depth (z).
2000
NiG5
100 0 --- -
... ....G
..
4..
. ... ..
S500-NG-
200
-
-
NG-NG2
100
300
350
400
450
500
550
600
Wavelength(nm)
Figure 3-5:Absorption depth of Schott neutral density glass filters
The only property necessary to calculate absorption coefficient using IPMS is the
speed of sound (c1). Although the speed of sound is usually known, it can be calculated
from Young's modulus, Poisson Ratio and density as illustrated in eq. (3.6). Table 3-1
54
summarizes these properties and the resulting speed of sound. The speed of sound of
glass is about 3.5 times larger than the speed of sound in water, the main constituent of
biological tissue.
Glass Type
Young's Modulus
2
Poisson Ratio
2
X 10~3 Kgm/(s mm )
Density
g/cm
3
Speed of
Sound (pm/ns)
NG1
62
0.229
2.49
5.37
NG10
62
0.229
2.47
5.39
NG9
62
0.229
2.44
5.42
NG3
62
0.229
2.44
5.42
NG4
62
0.227
2.43
5.43
NG5
62
0.229
2.43
5.43
Table 3-1 :Material properties and calculated speed of sound for several Schott neutral density
filters
3.3.3 Results
A sample displacement trace for glass filter NG1 is shown in figure 3-6. A single
pulse of Nd:YAG pump laser is absorbed by the glass sample at 0 ns, generating a
thermoelastic stress within the glass sample. The thermoelastic stress propagates through
the glass, reflecting off the free surface and causing a non-linear surface displacement
that is a function of the sample's absorption depth D and speed of sound c,. Before the
laser pulse is absorbed, the glass surface is stationary, and there is no surface
displacement. For times longer than 300 ns, shear waves influence the surface profile,
55
and the laser-target geometry is no longer one dimensional. The surface displacement
will eventually decay back to the zero baseline via thermal diffusion. The time scale for
diffusion is
Dc2/cX where a is the thermal diffusivity of the target. For the absorption
-
depths used in these experiments, the thermal diffusion time ranges from 0.02 - 6 s. The
maximum surface displacement is 45 nm and is a function of material properties and
incident laser energy only. This information could be used as an additional diagnostic
metric, although the need for an accurate local fluence may limit its potential.
50
E =550 mJ/cm
40
30
S
U
--- -- -- - - --
20
10
Laser pulse
0
-150
-75
75
0
150
Time (ns)
Figure 3-6:Sample displacement trace for Schott glass type NGl.
56
225
300
A Gauss-Newton non-linear least squares fitting algorithm was used to compute
the absorption depth and maximum displacement of each trace. Figure 3-7 depicts the
surface displacement with the best fit solution.
50
E =550mJ/cm 2
Df, =101 pm
40
--
- ----
- - --
30
S
Q
------ - -
20
ci~
----- -
10
- - - --
Laser pulse
y .....
0
-150
-75
0
75
150
225
300
Time (ns)
Figure 3-7:Sample displacement trace and best fit solution for Schott glass type NGI
The best fit surface displacement solution had an absorption depth of 101 Pm and
clearly matched the measured surface displacement. We repeated these measurements for
laser energies ranging from 12 - 440 mJ/cm 2 . The displacement traces and best fit curves
are shown in figure 3-8. The maximum displacement at the lowest fluence of 12 mJ/cm 2
displacement trace was 0.8 nm
57
35
Dft average = 102
pm
30
E =440 mJ/cm 2
25
-
E = 250
...
...
.....
20
--
J/--2
15
10
5
E
-
Laser pulse
105 mJ/cm
-
E
2
- --- --
12 m/
-----.
-.
-..
- ... ... ...
0--150
=
-100
-50
0
50
100
150
200
250
Time (ns)
Figure 3-8:Sample displacement traces and best fit solutions for Schott glass type NGL. Laser
fluence ranged from 12 - 440 mJ/cm 2 .
Figure 3-9 shows surface displacement traces with best fit solutions for all the
Schott glass filters used. The laser fluence was 550 mJ/cm 2 . For convenience, the
maximum displacements have been normalized.
58
D= 101 pm
NGI
- ----
0.8 -- -
0
- -
-Ci)
....-0.4
z
7 0p
-- -
--- --
-- --
D= 1750m
NG 5
------
0.2 Laser.. pu........e....
E
0
-75
0
75
150
~550 mJ/cm 2
225
300
Time (ns)
Figure 3-9:Sample displacement traces with best fit solutions for all the Schott glass filters
The best fit solutions perfectly match the measured surface displacement traces
for all glass samples. Table 3.2 summarizes the measured absorption depths for all glass
samples, compared to the values computed from transmission curves given by Schott
Glass Technologies. The measured absorption depth is given as the mean ± one standard
deviation.
59
Glass Type
Measured Absorption
Schott Given
(. = 532 nm)
Depth (60 points)
Absorption Depth
NG1
101 ±0.8 pm
106± 50 pm
NG10
187 ±1.5 pm
188± 11 pm
NG9
305 ±2.2 pm
306± 50 pm
NG3
435 ±3.5 pm
434± 40 pm
NG4
858
10 pm
856± 70 pm
NG5
1780
80pm
1790± 170pm
Table 3-2:Measured absorption depths for Schott neutral density filters compared with values
obtained from Schott Glass Technologies.
The IPMS technique yields absorption depths well within the range given by
Schott Technologies. The relatively large ranges given by Schott Glass Technologies,
indicate the inherent inaccuracies of extracting absorption depth of highly absorbing
materials using standard transmission spectroscopy. Since IPMS is not performed in
transmission mode, it may be a more accurate technique for measuring absorption depth
in highly absorbing materials.
To evaluate the consistency of IPMS, we calculated the absorption depth of NG 9
from 350 separate surface displacement traces. The traces were obtained over a period of
two months under different operating and alignment conditions. Figure 3-10 plots all 350
traces and clearly illustrates the robustness of IPMS. The y-axis limits correspond to the
absorption depth range given by Schott Glass Technologies.
60
350
Daverae
S
340
D
Q
320--
260-)
50 pm
306
--
310
280 -oe
XX)X
30
= 305.6 ±1.7 pm
- - XX
--
---. . .---
.
Xx
X
5
X
$
N
'5X2
0CM
'
-Y)XX
-
260
0
50
200
150
100
250
300
350
Data Traces
Vn
3.4~~~YTisXPatm PM4
Figure 3- 10:Calculated absorption depth for 350 traces obtained under various laser operating and
alignment conditions. The axis limits correspond to the absorption depth range given by Schott
Glass Technologies.
3.4 IPMS on Tissue Phantoms
3.4.1 Overview
IPMS is not limited to smooth specularly reflecting targets, and can be applied to
diffuse targets such as tissue. When measuring surface displacements on diffuse samples,
the interferometric system must include the beam expander and target focusing lens. This
ensures sufficient coherent light is reflected from the diffuse surface. To demonstrate
IPMS on a diffuse surface, the surface displacement of gelatin phantoms doped with an
61
absorbing dye were investigated. The absorbing dye was used to increase the absorption
of the tissue phantom to ensure one dimensional laser-tissue geometry. Potassium
Chromate (K 2 CrO 4 , Aldrich Chemical Company, Milwaukee, WI) was chosen as the
absorbing dye, since the absorption properties have been measured and it is reported to be
photochemically stable [55, 67]. These investigators estimated the absorption depth of a
35g/L K2 CrO 4 solution to be 10pm. The extinction coefficient (e), a more common
descriptor for absorbing dyes is given by the following expression:
E
(3.25)
1
ln(lO)C,,D
where C., is the molar concentration, and D is the absorption depth. The extinction
coefficient was calculated to be 2410 Lmol'cm'.
3.4.2 Experimental Setup
The experimental setup was similar to the setup used in the IPMS verification
experiments described in section 3.2, except the beam expander and target focusing lens
was used. The Q-switched Nd:YAG pump laser was operated at its third harmonic with a
wavelength of 354.7 nm. The spatial profile resembled a top hat with a FWHM diameter
of 6 mm. The temporal profile was Gaussian in shape with a FWHM pulse duration of
about 8 ns. The spatial and temporal profiles were similar to the profiles when the laser
was operated at 532 nm in the previous section. The laser fluences used in these
experiments were 70 mJ/cm 2.
62
Two separate batches (A and B) of potassium chromate solutions were prepared
with molar concentrations of 0.0188M, 0.0169M and 0.0154M, with estimated absorption
depths of 95 pm, 106 pm and 117pm respectively. These solutions were prepared so we
could verify the extinction coefficient of potassium chromate, and confirm the accuracy
of IPMS by differentiating subtle changes in absorption coefficient. To minimize
preparation errors, all potassium chromate solutions were filtered dilutions of two
0.169M stock solutions. The filters used were 0.2 pm sterile Acrodiscs (Gelman
Sciences, Ann Arbor, MI). All solutions were prepared with distilled water.
Tissue phantoms were prepared by mixing 10% of gelatin (G-1890 Type A gelatin,
Sigma, St. Louis, MO) by weight with a 0.0169M potassium chromate solution from
batch A. This enabled the influence of gelatin to be directly observed.
3.4.3 Results
Potassium Chromate Solutions
A sample displacement trace and best fit solution for a 0.0169M potassium
chromate solution from batch A is shown in figure 3-11. The best fit solution clearly
matched the measured displacement trace and had an absorption depth of 86 pm, which
was lower than the estimated value of 106 pm.
63
35
E" =70 mJ/cm 2
30- - D ,,= 86 pm--
- - - ... ..
...
25-
4S
20
--
--- ----- - -- - - - - - -- - - -
15
10
Laser pulse
5
-O
-
--- -- - - -
0
I
-20 0
0
I
200
f
400
I
600
800
Time (ns)
Figure 3-1 1:Sample displacement trace and best fit solution for a 0.0 169M potassium chromate
solution from batch A.
Table 3.3 summarizes the results for the three potassium chromate solutions from
batch A and B.
Molar Concentration
(A) Measured Absorption
(B) Measured Absorption
Estimated
(k = 354.7 nm)
Depth (30 points)
Depth (30 points)
Absorption Depth
0.0188M
75 ± 1.9 pm
86
0.0169M
86 ± 2.5 pm
96.3
0.0154M
93.2 ± 3.0 pm
106
2.0 pm
±
95
50 pm
1.8 pm
106~ 50 pm
2.4 pm
117 ± ~50 pm
Table 3-3:Measured absorption depth for three potassium chromate solutions from two different
batches (A and B) compared with estimated values.
64
These results show excellent agreement within each batch, clearly differentiating
subtle changes in absorption depth. However, there is some discrepancy between each
batch and the estimated absorption depth. These discrepancies are most likely due to
differences in preparing the potassium chromate solutions. If we calculate the extinction
coefficient based on the measured absorption depths, we get 2972 LmolFcm' and 2623
Lmol'cm' for batch A and B respectively. Both these values are reasonably close to the
value estimated from the literature (2410 Lmol'cm1 ), and the value we obtained via
standard transmission experiments (2800 Lmol'cm').
Our results highlight the uncertainty in obtaining accurate absorption depth
information from highly absorbing samples. In such cases, IPMS appears to be a very
suitable technique capable of detecting small absorption depth differences.
Gelatin Phantoms
A sample displacement trace and best fit solution for a gelatin doped tissue
phantom is shown in figure 3-12. The best fit solution clearly matched the measured
displacement trace and had an absorption depth of 136 pm. This absorption depth was
higher than the absorption depth of the 0.0169M potassium chromate solution alone (A 86 pm).
65
40
E =70 mJ/cm2
Df= 136 pm
35
30
25
S
Q
20
15
--
-- -- -----
-.
...
10
Laser pulse
5
-.
........
01
-2 0
0
200
400
600
800
Time (ns)
Figure 3-12:Sample displacement trace and best fit solution for a gelatin doped tissue phantom.
The tissue phantom was prepared with the 0.0 169M potassium chromate solution from batch A.
The absorption depth of gelatin doped tissue phantoms appeared to decrease with
repeated laser pules. Although this decrease was small, it was clearly detectable. Figure
3-13 compares the calculated absorption depth for the 0.0169M potassium chromate
solution with the gelatin phantom prepared with the same solution.
66
160
-- - - - - - - - -- - - - - - - - ++
150
140
-J
130
gelatin and
K 2 CrO 4 solution
120
K 2CrO 4 solution
-
110
100
-
U
90
-
80
-0
5
-------10
15
20
25
30
35
40
Data Points
Figure 3-13:Calculated absorption depth for 0.0169M potassium chromate solution (batch A) and
calculated absorption depth for a gelatin tissue phantom prepared with the same solution
It was unclear why the absorption depth decreased with each additional laser
pulse. The experiments were also performed on other gelatin types, including type A (G-
2625, Sigma, St. Louis, MO), type B (G-9391, Sigma, St. Louis, MO) and Knox
unflavored Gelatine (Nabisco, Hanover, NJ). The calculated absorption depth of these
phantoms exhibited dependence on gelatin type, concentration of potassium chromate
solution and laser fluence. Potassium chromate solutions alone were stable, although they
did exhibit slight photobleaching effects (< 8%) at high fluences. The absorption of all
gelatin type phantoms prepared without the dye was small (D ~ 3mm - 8 mm). The
specific role of gelatin type, potassium chromate concentration and laser fluence on
67
absorption depth was not investigated. Explanations for these effects may include
photochemical processes, or laser induced changes in spatial absorption properties of the
phantom. For example, the concentration of dye within the gel may be slightly altered
after each laser pulse, thereby creating inhomogeneous absorption. We have observed
these phenomena in gelatin phantoms prepared with India ink, and pumped with infrared
wavelengths where no photochemical effects play a role.
Gelatin type, absorbing dye and laser energy do influence optical properties of
tissue phantoms. Although some of these effects may be small, they were clearly
detectable with IPMS. We feel it is important to keep these effects in mind when
preparing tissue phantoms to be used for testing new techniques or applications.
3.5 Discussion
Interferometric Photomechanical Spectroscopy (IPMS) is a powerful non-contact
technique which accurately extracts the absorption depth of a sample. The absorption
depth is calculated from the sample's surface expansion after absorption of a short laser
pulse. The time dependent surface expansion is a function of optical penetration depth.
We have verified the accuracy of IPMS by calculating the absorption depth of
well characterized absorbing glass samples. The range of computed absorption depths
were more than an order of magnitude lower than the values given by the glass
manufacturer. In addition, IPMS was very robust, producing consistent results when
performed under different alignment and laser operating conditions. The consistent
accurate results of IPMS are primarily due to the high spatial and temporal resolution of
the interferometric system. IPMS relies on non-linear fitting to the time dependent
68
surface displacement of a sample. If the surface displacement is acquired with negligible
noise, or very high spatial resolution, the range of possible fitting parameters is narrow.
This is particular beneficial when fitting to only a portion of a sample's surface
displacement, as is the case for samples from long absorption depths. This is an attractive
feature which avoids the need for using the full three dimensional numerical wave
solution to extract absorption depths for weaker absorbing samples. The maximum
surface displacement could also be used to extract additional diagnostic information,
since it is a function of thermal expansion coefficient, Poisson ratio, density and specific
heat.
The challenge of diagnostic spectroscopy is to measure changes in optical
properties such as effective absorption depth, and equate these changes with specific
changes in tissue state. Often, changes in the optical properties are subtle, and must
therefore be measured with high resolution techniques. IPMS is a high resolution
technique capable of differentiating small changes in absorption depth. Further, IPMS
could be implemented into an optical fiber system and be used as a clinically valuable
minimally invasive diagnostic tool.
69
Chapter 4
Interferometric Photomechanical
Tomography (IPMT)
4.1 Introduction
There is great demand for non-invasive techniques to image deep sub-surface
absorbers such as tumors and blood vessels in highly scattering tissue. Such techniques
would be very useful for breast cancer screening, laparoscopic and endoscopic
procedures, and a variety of laser-tissue interactions including treatment of port wine
stains. The major challenge in imaging absorbers deep in tissue light is to maintain high
resolution despite significant tissue scattering. Coherent imaging techniques such as
optical coherence tomography and confocal microscopy cannot be used to detect deep
absorbers in tissue. Standard confocal microscopy has a penetration depth of only 100200 pm, and research based confocal microscopy have obtained depths up to 400 Pm [13,
19, 33, 77, 84]. Optical coherence tomography tends to have a slightly larger penetration
of about 1 -2 mm for highly scattering tissue [13, 84], although in our experience, a high
resolution clinical OCT instrument had a penetration depth of under less than 1 mm in
71
skin. The same instrument used in the gastrointestinal (GI) tract had penetration depths of
up to 2 mm [12].
There are other optical techniques that use diffuse light instead of coherent light
for imaging or characterization of sub-surface absorbers such as photon migration [99].
In addition, there are non-optical imaging techniques that are used for detecting deep
tumors. These techniques include X-ray radiography, magnetic resonance imaging and
ultrasound imaging. Although these techniques do not suffer from optical scattering, the
major limitation is usually associated with low contrast leading to insufficient sensitivity
for the detection for small tumors or blood vessels. Photoacoustic tomography
(optoacoustic imaging) is a new technique which utilizes high contrast optical absorption
to generate mechanical transients which propagate relatively unaffected through large
tissue sections (e.g. [26, 41, 55, 68, 101]). Photoacoustic imaging uses a pump probe
technique, where a short pulse of laser light is preferentially absorbed by a sub-surface
target such as a tumor or blood vessel. The sub-surface target expands and creates an
acoustic or stress wave that propagates to the tissue surface where it is measured by one
or a series of pressure transducers. Photoacoustic tomography essentially captures the
most beneficial attributes of optical and ultrasound imaging techniques by combining
high optical contrast with deep penetration properties of sound transients.
In general, non-invasive photoacoustic medical imaging techniques must be
performed in an epitaxial or so called backward mode. Consequently, standard front
surface pressure transducers cannot be used as they would physically block the incident
laser light. To solve this problem, investigators have used a variety of creative
techniques, including piezoelectric transducers with separated light and sound fields,
72
transparent transducers measuring optical reflectance, annular piezoelectric elements, and
acoustic conductors where sound generated by an obliquely incident laser pulse
propagates predominantly normal to the tissue surface [8, 9, 17, 55, 57, 69].
Alternatively, an interferometric technique that is inherently epitaxial could be used to
image sub-surfaces absorbers. The ultra high spatial and temporal resolution
interferometer discussed in Chapter 2 is ideal for imaging sub-surface absorbers by
measuring surface di-splacements caused by acoustic transients. We call this technique
Interferometric Photomechanical Tomography. We use the term photomechanical rather
than photoacoustic or optoacoustic, because we measure surface deformation and not
acoustic transients.
In addition to being epitaxial, there are many significant advantages for using an
interferometric surface monitoring technique. First, interferometric monitoring of surface
displacement is a non-contact method, which is highly beneficial in medical applications.
Second, IPMT is a point measurement, limited only by the numerical aperture of the
imaging optics. It has been shown that the lateral resolution for optoacoustic imaging is
limited by transducer size [41]. Third, the interferometer has a very high bandwidth
limited only by the digitizing system (500 MHz) and does not suffer from calibration
uncertainty. Fourth, the interferometric system could easily be implemented into an
optical fiber setup, enabling minimally invasive sub-surface imaging during catheter
based procedures such as endoscopy. Fiber optic components and related technologies for
minimally invasive interferometric techniques such as optical coherence tomography
already exist and could be readily adapted to our interferometric system. Lastly, the
interferometric system could be modified to produce an array of probe beams to allow
73
simultaneous point measurements. This would enable full three-dimensional imaging in a
single shot
Although interferometric surface monitoring of photomechanical transients has
many significant advantages over measurements made by pressure transducers, it has not
been pursued in the medical community. In fact, we only found one article that used an
interferometric system for medical imaging of acoustic transients [52]. This is surprising
because the pioneering studies by Albagli and coworkers showed great potential [2, 3,
48]. The reason why interferometric techniques were not pursued for medical imaging
may be due to the relatively low spatial resolution (- ± 4 nm) reported in these earlier
studies. Photomechanical transients from deep absorbers are attenuated inversely to
distance traveled, and consequently sensitive surface displacement or stress detection is
necessary. The interferometric system described in Chapter 2 is capable of a spatial
resolution which is 50 times more sensitive than the interferometric system in these
earlier studies. We feel that this significant increase in spatial resolution, combined with
the advantages listed above, should enable interferometric surface monitoring techniques
such as IPMT to become a powerful tool for imaging in vivo deep sub-surface absorbers
such as tumors and blood vessels. In addition, current work in photoacoustic image
reconstruction for stress based techniques can be applied to interferometric
photomechanical techniques.
In this chapter, we have successfully applied Interferometric Photomechanical
Tomography (IPMT) to image sub-surface blood vessels in a phantom model and in vivo.
We used a simple scanning technique to extract vessel size and depth information.
74
4.2 Verification of Interferometric Photomechanical Tomography
4.2.1 Overview
IPMT was performed in several well characterized tissue-blood vessel phantom
models, and in the forearm of a volunteer human subject. The tissue phantoms consisted
of either pure water, 1% Intralipid solution or 7% Intralipid solution. Although the
majority of investigators use 1% Intralipid solutions as tissue phantoms, we feel a 7%
Intralipid solution more accurately mimics highly scattering biological tissue. The blood
vessel phantoms consisted of different diameter polyimide tubes. The polyimide walls
were acoustically thin (19 pm - 50 pm), so stress reflections due to impedance mismatch
between the vessel wall and surrounding medium were not observed. Stress reflections
were observed in thicker walled phantom vessels. Although stress reflections could be
used to extract size information, such phantom vessels are not realistic.
Currently, photoacoustic imaging is performed by measuring acoustic transients
generated from a sub-surface absorber at multiple surface locations. Each spatially
resolved stress history is combined in a reconstruction algorithm to create an image of the
sub-surface absorber. An excellent example of this technique was recently illustrated by
Hoelen and co-workers who imaged phantom blood vessels using a transducer array in
the forward mode [23, 41, 42]. IPMT could also be performed in a similar fashion, where
an array of interferometer probes measuring surface displacements are used to reconstruct
sub-surface tomography. However, we used an alternate technique which has not been
75
previously used in photoacoustic imaging. It is a simple scanning technique which avoids
the need for complex image reconstruction techniques.
The scanning technique is illustrated in figure 4-1. The interferometer probe beam
is centered within the incident pump beam on the target surface. Single shot
measurements were obtained across the sample's surface by moving the sample via a
digital micrometer stage. Each spatially resolved time trace was stored and combined
later to form a tomographic image.
Incident
laser pulse
Interferometer
probe beam
Phantom scan
Tissue
phantom
Figure 4-1:IPMT scanning setup. Surface displacement traces across the sample were obtained by
scanning the tissue phantom on a digital micrometer.
76
4.2.2 Experimental Setup
Laser System
The pump laser used was a Q-switched Nd:YAG laser, operating at its
fundamental wavelength of 1064 nm. This wavelength was chosen since it is absorbed
more strongly by blood than the surrounding tissue and therefore could also be used for
the in vivo experiments described later. The pulse duration of the laser beam was about 8
ns, and had a Gaussian temporal shape. The spatial profile was analyzed using a CCD
Camera (Cohu), and found to resemble a flat top. The full width half max (FWHM)
diameter of the raw beam was measured to be about 6.6 mm. For the scanning
experiments, the beam was focused to a spot with a diameter of 0.6 mm. The lens had a
long focal length of 1000 mm to ensure the beam diameter did not vary significantly at
the target. Figure 4-2 illustrates the spatial energy distribution of the focused spot at the
target.
The Q-switched Nd:YAG was directed towards the sample by gold mirrors. The
angle of incidence at the sample was about 15'. A diagram of the setup is shown in figure
4-3. The laser fluence used for the experiments ranged from 0.2 J/cm
depending on the amount of scattering of the tissue phantom being probed.
77
2
to 1.5 J/cm 2
3
2.5
S1.5
0.5
0
Distance (mm)
Figure 4-2:Spatial profile of a focused Q-switched Nd:YAG laser beam at 1064 nm. The FWHM
diameter of the spot was about 0.6 mm
78
Interferometric
system
Gold
mirror
Interferometer
w probe beam
Nd:YAG
1064 nm
850
Phantom model
Figure 4-3:Schematic of experimental setup for IPMT
Phantom Model
The tissue phantoms consisted of either pure water, 1 % Intralipid, or 7%
Intralipid solution. The Intralipid solutions were made by diluting stock 10% Intralipid
solution (Pharmacia Inc., Clayton, NC)). The optical properties of Intralipid at 1064 nm
were obtained from Royston and coworkers [80]. Note, the authors quote the optical
properties of Intralipid as a percentage of 10% stock solution, not the total percent. The
reduced optical scattering coefficient at 1064 nm were is 0.65 mm-' and 4.55 mm-' for 1%
and 7% Intralipid solution respectively. A 7% Intralipid solution more accurately mimics
highly scattering tissue such as skin, which has a reduced scattering coefficient of 3.55
mm-' [18].
79
The blood vessel phantoms consisted of thin walled polyimide tubes of diameters
ranging from 200 - 1000 pm (Cole-Parmer, Vernon Hills, IL). The wall thickness was
between 19 pm - 50 pm and were considered acoustically thin. The polyimide tubes did
not absorb significant laser energy. A dilute solution of black India ink (Higgins,
Bellwood IL) was used to simulate the absorption properties of blood. The absorption
coefficient of whole blood at 1064 nm is about 0.5 mm' [76, 94]. The optical properties
of India ink at 1064 nm were obtained from Royston and coworkers and verified using
IPMS [80]. A 0.125% India ink solution (by volume) has an absorption coefficient of
0.45 mm-1 or absorption depth of 2.2 mm and was used in all experiments. A simple
syringe setup was used to flow the dilute India ink solution through the polyimide tubes
after each measurement to prevent any ink particle separation.
In Vivo Human forearm
A vessel in the forearm of a human volunteer was also imaged using IPMT. It was
not feasible to accurately scan the forearm through the pump and laser beam. Instead,
single measurements were made on spatial locations either on or off the vessel. The
vessel location was determined visually, by the presence of a faint blue/green line. A thin
layer of water was used on top of the skin surface. This was not to used to generate a
specular surface, but used simply to aid in alignment. This would not be necessary for a
fiber optic based clinical system. Since there was no scanning, the 1064 nm Q-switched
Nd:YAG laser was not focused onto the surface of the skin. We also used a clinical OCT
system [11] to image the same location.
80
4.2.3 Results
Pure Water
A sample surface displacement trace of a vessel tissue phantom is shown in figure
4-4. The surface displacement is caused by the absorption and subsequent expansion of a
sub-surface 1000 pm phantom blood vessel and surrounding pure water. The incident
pump laser and interferometric probe beam were centered directly above the phantom
blood vessel. A single pulse of Nd:YAG pump laser is absorbed by both the sub-surface
absorber and surrounding water at 0 ns. The initial surface displacement is due to the
absorption of laser radiation in the surrounding Intralipid. The surface displacement
caused by the submerged phantom blood vessel starts at 2.52 ps. The time delay
represents the time for the acoustic wave generated by the phantom vessel to reach the
surface. The vessel depth is calculated by multiplying the delay time with the speed of
sound in pure water. The longitudinal speed of sound in pure water at 25'C is 1.496
mm/ps [16]. The vessel depth was therefore calculated to be 3.78 mm. The upper axis of
figure 4-4 represents depth as computed from delay time. The actual vessel depth was
3.82 mm, which is 40 pm less than the measured value. The actual depth is obtained by
comparing the height of the vessel surface before any water is added, to the height of the
water surface after the water is added. This is achieved by using a digital micrometer to
move the phantom surface (vessel or water) vertically into the focal plane of the
interferometric probe. The vertical resolution of this technique is about 10 pm, and is a
function of the micrometer and the depth of focus of the interferometer probe beam.
This technique can be used to measure the speed of sound in an unknown liquid
by measuring transit time of an acoustic wave through a known distance. We tested the
81
accuracy of this technique by calculating the speed of sound of a planar wave in pure
water. The planar wave was generated by irradiating a Schott glass filter (NG1) through
pure water. The average speed of sound at ten different depths was measured to be 1.493
- 1.513 mm/ps, which closely matches the given value of 1.496 mm/ps for pure water.
Alternatively, assuming the speed of sound is 1.496 mm/ps , the average depth error was
14 pm. We repeated this measurement for 7 % Intralipid solution, and measured the
speed of sound to be 1.487 - 1.529 mm/ps, which is slightly higher than for pure water.
This value is slightly lower than the value reported for human milk, which is 1.54 mm/ps
[34]. Although these differences are small they should be taken into account when
extracting accurate depth information.
82
Depth (mm)
45 -
0
1.5
3.0
--- -- -
--
4.5
7.5
6
9.0
10.1
12
40
35
S30
-- - - -
- - - --- - - -
*-25
0 20
-
----------
--
- -
--------- ----- - ---
4
5
--
15
10
-lI
Laser pulse
0
1
2
3
Time (ps)
6
7
8
Figure 4-4:Sample surface displacement of a vessel phantom model consisting of a 1000 pm
polyimide tube surrounded by pure water.
Figure 4-5 illustrates displacement traces obtained from the same 1000 Pm
phantom vessel at different depths. The maximum displacement decreases linearly with
depth as expected. The calculated vessel depths were 3.78 mm, 5.79 mm, 7.24 mm,
which were within ± 45pm of the actual depths (measured via digital micrometer).
Although the depth errors were small, they were significantly larger than the depth errors
for the planar wave experiments (14 pm). The larger depth errors occurred because wave
propagation effects for non-planar waves such as diffraction were not taken into account.
Wave propagation effects should be accounted for if accurate depth information is
necessary. Nonetheless, the high spatial resolution of the interferometer is beneficial
83
when measuring surface displacement caused by acoustic waves generated from deep
absorbers.
Depth (mm)
45
0.
1.5
3.0
4.5
6
7.5
9.0
10.1
12
40
35
-- - --- --- ---- --- - - -
-... .-
---- ----- - - -
30
4-25
20
15
10
- -.
0
-
----
. . .. . . . .. .
5
Laserpul0
1
2
3
4
Time (ps)
5
6
7
8
Figure 4-5:Sample surface displacement traces of a 1000 pm phantom vessel at different depths.
The surrounding medium was pure water.
Surface displacement traces obtained at different spatial locations across the
phantom model can be combined to form a tomographic image of the time dependent
deformation. No other processing is necessary. A sample tomographic image for a 1.53
mm deep 495 pm diameter phantom blood vessel is shown in figure 4-6. A contour plot
is shown in figure 4-7. The x-axis is the scanning axis, and the y-axis is the temporal or
depth axis.
84
30
~25
-----
10
---
3
15-00
Time (ps)
0
-0.5
Distance (mm)
Figure 4-6:Sample tomographic image of a 495 pm diameter sub-surface phantom vessel in pure
water. Vessel depth was 1.53 mm.
85
3
4.5
2.5
3.8
2
3.0
1.52.
1.5
0.5
0.8
0
0
-0.4
-0.2
0
0.2
0.4
0.6
Distance (mm)
Figure 4-7:Sample contour image of a 495 pm diameter sub-surface phantom vessel in pure
water. Vessel depth was 1.53 mm.
Figure 4-8 compares tomographic images for three different phantom vessels with
diameters of 1000 pm, 495 pm and 198 pm respectively. The vessel depths were all
about 1.5 mm. The bar in figure 4-8, and all other figures represents 500 pm. The vessel
sizes are clearly distinguishable both in horizontal scanning axis (x-axis), as well as in the
temporal or depth axis (y-axis).
86
2.5
4
3.5
1.5
215
0.5.1
d=495 umn
d=00Oumn
d=198um
0.5
Figure 4-8:Sample contour images of three different sized phantom vessels in pure water, The
depths of the phantom vessel are all about 1.5 mm.
One way to obtain a quantitative measure on vessel size is to plot the maximum
surface displacement as a function of length or scanning distance (x-axis) as is illustrated
in figure 4-9. The maximum displacement profiles have been normalized and fitted to a
Gaussian function. The width of the three profiles calculated at 60% of the maximum
displacement are 970 pm, 510 pm and 230 pm for phantom vessel diameters 1000 pm,
495 pm and 198 pm respectively. The maximum displacement profile is clearly a good
indicator of vessel size in a non scattering medium.
87
I
I
i
I
I
-
-
0.9 -- - --
S 0.6
I
---
0.5
04
-
d=10OO pm
~d
=495
S 0.1 --- ----1.5
---1
-
-d=198pm
-0.5
0
0.5
1
1.5
Length (mm)
Figure 4-9:Normalized maximum displacement profiles for three different phantom vessels in
water. The diameters were of the vessels were 1000 pm, 495 pm and 198 pm. Gaussian functions
have been fitted to the profiles.
Comparison of Pure Water, 1 % Intralipid Solution and 7% Intralipid Solution
To access the effect of scattering, we repeated these experiments in 1% Intralipid
solution and 7% Intralipid solution. Figure 4-10 compares tomographic images for a 1000
pm phantom vessel in pure water, 1% and 7% Intralipid solution. The depth in all three
cases was about 1.5 mm. The vessel images for the 1% and 7% Intralipid solutions have
been broadened in the horizontal scanning axis (x-axis), but not the temporal or depth
axis (y-axis). The temporal profile is not strongly affected by the scattering properties of
the phantom medium.
88
4.5
3
4
2..53.5
1.5
0.
Pur wate
intralinid
1%
I
0
0.
~5
Figure 4-1 0:Sample tomographic images for a 1000 pm phantom vessel in pure water,
1% and
7% Intralipid solution. The depth in each media was about 1.5 mm.
The normalized maximum displacement profiles are shown in Figure 4-11. The
width of the three profiles calculated at 60% of the maximum displacement are 970 Pim,
1250 pm and 1720 pm for a 1000 pm diameter vessel in pure water, 1%
and 7%
Intralipid solution respectively. The maximum displacement profiles are broadened by
scattering. The 1%
Intralipid image and the pure water image are relatively similar
indicating that scattering in 1% Intralipid solution is not very significant.
89
0.9
-.
0.7 --
-
- -
--
- - -
-- --- --- -
.-
-
-----0.6 - ----------. .. .- --- .----. . .. . ...
0 .5 -.. . .. . .--0 .4
---
- ---- -- -
- - -- -- --
- --
--
0.8 - ---- --
--
--
-
- - - - --
-
.
.
- -
-
--
- - --
-
-
. .. -.-.
--
-
- -
7%
-
0.3
0.2
1%
-
0.
-3
-
-
--
water
-2
-1
0
1
2
3
Distance (mm)
Figure 4-1 1:Normalized maximum displacement profiles for a 1000 Pm diameter vessel in pure
water,
1% and 7% Intralipid solution. Gaussian functions have been fitted to the profiles.
7% Intralipid Solutions
A sample surface displacement of a 300 pm vessel tissue phantom is shown figure
4-12. The incident pump laser and interferometric probe beam were centered directly
above the phantom blood vessel. A single pulse of the Nd:YAG pump laser is absorbed
by both the sub-surface absorber and surrounding Intralipid at 0 ns. The initial surface
displacement is due to the absorption of laser irradiation in the surrounding Intralipid.
Since the Intralipid absorbs more laser energy than pure water, the initial surface
displacement is greater than it was for the pure water phantoms (see figure 4-4). The
surface displacement caused by the submerged phantom blood vessel starts at 0.67 ps.
90
The time delay represents the time for the acoustic wave generated by the phantom vessel
to reach the surface. The vessel depth is calculated by multiplying the delay time with the
speed of sound in 7% Intralipid, which was measured to be 1.508 mnpm (see the results
section under Pure Water). The vessel depth was therefore calculated to be 1.01 mm. The
upper axis of figure 4-12 represents depth as computed from delay time. The actual
vessel depth was 1.04 mm, which is 30 pm more than the measured value. The actual
depth was obtained by comparing the height of the vessel surface before any Intralipid is
added, to the height of the Intralipid surface after the Intralipid is added (as mentioned
before).
Depth (mm)
0
.8
1.5
2.3
3.0
3.8
4.5
5.3
60
50
40
S
30
07
4J
20
10
Laser pulse
------ ----------------------------------0.
-0.5
0
0.5
1
1.5
2
2.5
3
3.5
Time (ps)
Figure 4-12:Sample surface displacement of a vessel phantom model consisting of a 300 Pm subsurface polyimide tube surrounded by 7 % Intralipid solution. The vessel depth was 1.04 mm.
91
Figure 4-13 illustrates displacement traces obtained for the same 300 Pm phantom
vessel at different depths. The calculated vessel depths were 1.04 mm, 1.65 mm, 2.13
mm, 2.72 mm, 3.27 mm, 3.82 mm which were within ± 60pm of the actual depths
(measured via digital micrometer). The maximum displacement no longer decreases
linearly with depth. This is because the amount laser energy reaching the deeper vessels
is significantly reduced due to scattering. In non-scattering media, the maximum
displacement was proportional to depth (see figure 4-5). For a highly scattering medium,
laser light penetration may be a limiting factor when imaging subsurface absorbers.
Depth (mm)
0
60 -
.8
2.3
1.5
3.8
4.5
5.3
-
40 --- - ---- -- -- - -- ---- ---
- - ----- - -
30
3.0
2 0 ---- -
----
--- - -- - -- --- - --
- --- - - _-----
---
-- - --
----
10-Lase pulse
0
-0.5
-
0
0.5
1
1.5
2
2.5
3
3.5
Time (ps)
Figure 4-13 :Sample surface displacement traces of vessel phantoms at different depths in a 7%
Intralipid solution. The phantom vessel size was 300 pm in diameter.
92
Surface displacement traces obtained at different spatial locations across the
phantom model were combined to form a tomographic image of the time dependent
deformation. A sample tomographic image for a 1.36 mm deep 1000 pm diameter
phantom blood vessel is shown in figure 4-14. A contour plot is shown in figure 4-15.
The x-axis is the scanning axis, and the y-axis is the temporal or depth axis. The initial
surface displacement due to laser absorption be the Intralipid is clearly visible.
120
10
3
3
200
-1
-3
Distance (mm)
Figure 4-14:Sample tomographic image of a 1000 pm diameter sub-surface phantom vessel in a
7% Intralipid solution. Vessel depth was 1.36 mm.
93
3
4.5
0.5
.8
0
0
-0.5
-2
-1
0
1
2
3
Distance (mm)
Figure 4-15:Sample contour image of a 1000 pm diameter sub-surface phantom vessel in a 7%
Intralipid solution. Vessel depth was 1.36 mm.
Figure 4-16 compares tomographic images for three different phantom vessels
with diameters of 1000 pm, 495 pm and 198 pm respectively. The vessel images are
significantly broadened in the horizontal scanning axis (x-axis). The temporal or depth
axis (y-axis) has not been significantly affected by the scattering properties of the
phantom medium.
94
05
Figure 4-1 6:Sample contour images of three different sized phantom vessels in a 7% Intralipid
solution. The depths of the phantom vessel are all about 1.4 mm.
The normalized maximum displacement profiles are shown in Figure 4-17. The
width of the three profiles calculated at 60% of the maximum displacement are 1720 pm,
1370 pm and 1150 pm for phantom vessel diameters 1000 pm, 495 pm and 198 pm
respectively. The maximum displacement profile is no longer a good indicator of vessel
size in a highly scattering medium. However, the temporal profile is still a good indicator
of vessel size, since it is not significantly affected by the highly scattering medium. This
clearly illustrates the advantage of using wave propagation over diffuse light propagation
for obtaining high resolution images in highly scattering mediums. To underscore this
point further, the temporal profile of three different phantom vessels with diameters of
1000 pm, 300 pm, and 198pm is shown in figure 4-18. The depth of all the three
phantoms was about 2.2 mm. The temporal profile and maximum displacement is
significantly different for the each phantom vessel and could both be used to extract size
information. The exact temporal profile is a function of vessel size, and wave propagation
properties including diffraction and attenuation.
95
S 0.8
-
-- -
1000pm
E 0.2
Z
0.1
-
-
. -
--- --.--
-
-
-- ----
-
-
-- - - ---
- - - -
- --
--
-- --
0.4 - --
--
-
---
--
-
0 .5 -
0.3 - -
-
--
------
---.6
-.
-
-
-x
-
0.9 --
4 9 5pm . . . . . ... . . . . .
. .. . .. . . ..
- 198pm
-
0 -0
-3
-2
-1
0
1
2
3
Distance (mm)
Figure 4-17:Normalized maximum displacement profiles for three different phantom vessels in a
7% Intralipid solution. The diameters were of the vessels were 1000 Pm, 495 pm and 198 Pm.
Gaussian functions have been fitted to the profiles.
96
Depth (mm)
0
45
.8
1.5
2.3
40
-----------
.
4.5
3.8
3
5.3
.......
- -------
- - - - --
1000pm
-
35
30
-
---
-
-
\300pm........... --
25 -..
S
20
-
---
--------
-- -
--
- - -- - - -
0
-
.198pm
15 10 - L a s er p
u ls e
- -
0
I.~5
-
-
5 ---
-------0
0.5
1
1.5
2
2.5
3
3.5
Time (ps)
Figure 4-18:Sample surface displacement traces of three different vessel at a depth of 2.2 mm in a
7% Intralipid solution. The diameters of the vessels were 1000 pm, 300 prm and 198 pm.
In Vivo human forearm
Since a scanning setup was not feasible, single measurements were made on
spatial locations either on or off the vessel. Figure 4-19 illustrates two sample surface
displacement traces caused by the absorption of 1064 nm laser radiation within the
forearm. The top trace was obtained when the pump and probe beam were centered on
top of the vessel, and the bottom trace was obtained when the pump and probe beam were
centered to the side of the vessel. In both cases, the alignment was not optimized, and the
baseline STD was 0.25 nm. A fiber optical setup would circumvent alignment
97
difficulties, and increase spatial resolution. Nonetheless, the presence of the blood vessel
is clearly detected.
The initial surface displacement is caused by absorption of laser radiation within the
water layer. At 1.6 ps, there is an increase in surface displacement owing to absorption
within the forearm skin. The increase of absorption is expected, since skin contains
additional chromophores such as melanin and blood. The water thickness is estimated
from the transit time of 1.6 ps to be 2.4 mm. The displacement in the top trace caused by
the submerged blood vessel starts at 2.2 ps. Using a speed of sound of 1.5 mm/ps for skin
[34], the vessel depth was estimated to be 0.9 mm. This appears to be a realistic estimate,
since the vessel was faintly visible from the surface.
98
Depth (mm)
70
0
1.5
3
60
4.5
--
6
7.5
9
- ............. ----- --
10.5
12
............
On vessel
50
.
. . . . ..
...
.
....
....
40
30
0
20 -- --- ----
O ff vessel
...-.
- - ---. . .... ...
Laser pulse
-.
10
...
-.
.....
.....
-.. .......................
0
-1
...........-
0
1
2
3
4
5
6
7
8
Time (ps)
Figure 4-19:Two sample surface displacement traces obtained from the forearm of a human
volunteer. The top trace was obtained directly above the vessel, and the bottom trace was
obtained to the side of the vessel. The vessel depth is estimated to be 0.9 mm below the skin
surface.
An OCT image was also taken directly on top of the same vessel and is shown in
figure 4-20. The bar indicates 500 pm. Although the image indicates some texture, there
is no clear indication of a blood vessel.
99
Figure 4-20:Sample OCT image taken directly on top of the same vessel in the forearm of a
human volunteer. There is no clear indication that there is a sub-surface blood vessel.
4.3 Discussion
In this chapter, we have developed and verified a new technique called
Interferometric Photomechanical Tomography (IPMT). IPMT was used to image subsurface phantom blood vessels as well as an in vivo blood vessel in a human forearm.
IPMT was performed in a simple scanning mode to obtain tomographic images of subsurface phantom vessels. The scanning technique was not feasible for the in vivo
experiments. A clinical optical fiber based interferometric system would be more suited
for in vivo applications.
We demonstrated IPMT in both non-scattering and highly scattering media
including real tissue. In non scattering media, the scanning technique produced accurate
tomographic images in both the horizontal or scanning axis, and the temporal or depth
axis. A quantitative measure of vessel size was achieved by comparing maximum surface
100
displacement profiles as a function of the scanning. The horizontal or scanning axis
resolution in a non-scattering medium is a function of the laser spot size.
For highly scattering media, the horizontal or scanning axis resolution is
significantly degraded. The maximum displacement profiles are broadened and can no
longer be used to extract vessel size. The temporal profile however, is still a good
indicator of vessel size, since it is not significantly affected by a scattering medium. This
result highlights the advantage of using wave propagation over diffuse light propagation
for obtaining high resolution images in highly scattering mediums. Phantom models
made with 1% Intralipid solution behaved more like water than the 7% Intralipid. We feel
a 7% Intralipid solution more accurately mimics highly scattering tissue, and should be
used in tissue phantom experiments.
The vessel depth in both non-scattering and scattering mediums were accurately
measured to within ± 60 pm. The depth error resulted because wave propagation effects
such as diffraction were not taken into account. For planar geometry, the depth error was
significantly lower (14 pm). Wave propagation effects should be accounted for if
accurate depth information is necessary.
Interferometric Photomechanical tomography is a potentially powerful noncontact technique used to image sub-surface absorbers in highly scattering media. A
simple scanning technique can be used to accurate extract the size and depth of a subsurface absorber. In highly scattering media, the size information must be extracted from
the temporal profile. A clinical scanning setup would enable sub-surface tomographic
images similar to ultrasound. IPMT could be combined with an image processing
algorithm to increase the horizontal resolution in scattering media. For example, a simple
101
algorithm could utilize the spatially dependent transit times to reduce spatial broadening.
IPMT could also be used with more complex reconstruction algorithms, similar to ones
currently used for photoacoustic (e.g.[41]). Since IPMT could be incorporated into an
optical fiber setup, it may have real clinical value.
102
Chapter 5
Conclusions and Future Work
5.1 Achievements of This Thesis
We have developed two novel minimally invasive diagnostic techniques. The first
technique is termed Interferometric Photomechanical Spectroscopy (IPMS) and measures
the effective optical absorption depth of a sample. We used IPMS to measure effective
absorption depth of both diffuse and speculary reflecting targets including well
characterized colored glass samples and gelatin based tissue phantoms. The second
technique is termed Interferometric Photomechanical Tomography and images subsurface absorbers in scattering media. We used IPMT to image sub-surface blood vessels
in a phantom model and in vivo. Both techniques measure surface deformation of a target
after absorption of a short laser pulse, and relate this surface deformation to the target's
spatially resolved optical, thermal and mechanical properties. We have also developed an
ultra-high resolution interferometric system to be used in both these techniques.
103
The specific achievements of this thesis are given in the following three sections:
Instrumentation, development and verification of Interferometric Photomechanical
Spectroscopy, development and verification of Interferometric Photomechanical
Tomography.
5.1.1 Instrumentation
We have designed and built an ultra-high resolution interferometric capable of
angstrom spatial resolution, and sub nanosecond temporal resolution. This configuration
was based on a previous version by Yablon and coworkers [108]. The spatial resolution
has been improved by more than an order of magnitude. The primary reason for this
improvement is the implementation of a dual balance detection system, as well as a
gradium index lens. The best spatial resolution achieved on a highly reflecting target was
0.08 nm. The best spatial resolution achieved on a gelatin tissue phantom was 0.16 nm.
The improvement in spatial resolution is useful for detecting weak surface
deformation when imaging deep sub-surface absorbers, or when differentiating subtle
changes in optical properties.
5.1.2 Development and Verification of InterferometricPhotomechanicalSpectroscopy
A novel minimally invasive spectroscopic technique called Interferometric
Photomechanical Spectroscopy (IPMS) was developed and verified. IPMS is a powerful
non-contact technique which accurately extracts the absorption depth of a sample. The
absorption depth is calculated from the sample's surface expansion after absorption of a
short laser pulse. The time dependent surface expansion is a function of optical
104
penetration depth. IPMS produced repeatable results with standard deviations of about 1
% of the sample's absorption depth. The mean extracted absorption depths were well
within the range given by the manufacturer.
5.1.3 Development and Verification of InterferometricPhotomechanicalTomography
A novel minimally invasive imaging technique called Interferometric Photomechanical
Tomography (IPMT) was developed. IPMT was implemented in a simple scanning setup
to image sub-surface phantom blood vessels in highly scattering media. IPMT was also
used to image a blood vessel in the forearm of a human volunteer.
We compared sub-surface images of non-scattering and highly scattering media.
The vessel depth in both non-scattering and scattering mediums were accurately
measured to within ± 60 pm, although this value can be reduced by including the effects
of wave propagation. The spatial or scanning resolution was significantly degraded when
scattering was present. The temporal profile was not significantly affected by the highly
scattering medium, and could be used to extract vessel size. This result highlights the
advantage of using wave propagation over diffuse light propagation for obtaining high
resolution images in highly scattering mediums.
We were able to clearly detect the presence of a blood vessel in vivo. The same
vessel was not visible using optical coherence tomography. Interferometric
Photomechanical tomography is a potentially powerful non-contact technique to image
sub-surface absorbers in highly scattering medium.
105
5.2 Future Research Directions
This thesis has illustrated two potentially useful diagnostic techniques, based on a
high resolution interferometric system. Although instrumentation advancements are
possible (as described in section 2.3), we feel future research should be directed toward
enabling real clinical applications. Four relevant areas are discussed below.
5.2.1 Comprehensive Comparison with Pressure Transducers
Interferometric techniques compare favorably against diagnostic techniques
which measure surface stress instead of surface displacement. This is primarily because
interferometry is a non-contact epitaxial method capable of high resolution point
measurements. However, a comprehensive comparison with state of the art pressure
transducers under similar conditions should be done, and clinical practicalities explored.
5.2.2 ClinicalInstrumentationDevelopment
The current interferometric system should be implemented into a optical fiber
based setup. This would enable both IPMS and IPMT to be used for in vivo applications.
Implementation of the current system into an optical fiber setup should not be difficult
since interferometric fiber based technologies already exist for other techniques like
optical coherence tomography. A scanning system or probe array could also be developed
enabling fast 3-dimensional imaging.
106
5.2.3 Image Reconstruction Development
Although a simple scanning technique can be used to obtain reasonable tomographic
images, image processing or reconstruction will improve spatial resolution in scattering
media. Reconstruction techniques can be as simple as utilizing the spatially dependent
transit times to reduce spatial broadening, or as complex as current photoacoustic
imaging reconstruction techniques (e.g. [41]).
5.2.4 Simple ClinicalApplications
Finally, simple clinical trials are necessary to confirm the diagnostic potential of
both IPMS and IPMS. For example, IPMS could be used to compare effective absorption
depth between pathological and non-pathological tissue. IPMT could be used during
endoscopic procedures to localize sub-surface blood vessels in the gastrointestinal (GI)
tract.
107
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