All-Optical Photoacoustic Imaging with Fiber-Optic Mach-Zehnder Interferometers / submitted by Johannes Bauer-Marschallinger

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JOHANNES KEPLER UNIVERSITY LINZ Altenberger Str. 69 4040 Linz, Austria www.jku.at Submitted by Johannes Bauer-Marschallinger Submitted at Institut für Mikroelektronik und Mikrosensorik Supervisor and First Examiner

Univ.- Prof. Dr. Bernhard Jakoby

Second Examiner Univ.- Prof. Dr. Günther Paltauf Co-Supervisor Dr. Thomas Berer Februar 2019

All-Optical

Photoacoustic Imaging

with Fiber-Optic

Mach-Zehnder

Interferometers

Doctoral Thesis

to obtain the academic degree of

Doktor der technischen Wissenschaften

in the Doctoral Program

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EIDESSTATTLICHE ERKLÄRUNG

Ich erkläre an Eides statt, dass ich die vorliegende Dissertation selbstständig und ohne fremde Hilfe verfasst, andere als die angegebenen Quellen und Hilfsmittel nicht benutzt bzw. die wörtlich oder sinngemäß entnommenen Stellen als solche kenntlich gemacht habe.

Die vorliegende Dissertation ist mit dem elektronisch übermittelten Textdokument identisch. Linz,

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Danksagung

Ich freue mich sehr, meine mehrjährige Forschungsarbeit nun auch in Form dieser Dissertation präsentieren zu können. Mein persönliches, berufliches und akademisches Umfeld hat mich dabei maßgeblich unterstützt. Vielen Menschen bin ich daher von ganzem Herzen dankbar. Zu allererst möchte ich mich natürlich bei meinen Eltern bedanken, die es mir ermöglichten und mich ermutigten eine höhere Schule und die Universität zu besuchen. Gerade in Österreich, wo Kinder oft keinen höheren Bildungsgrad wie ihre Eltern erreichen, ist es eine besondere Leistung von Fritz und Christine, dass all ihre vier Kinder eine höhere Schule erfolgreich besucht haben und drei ein Doktorratsstudium absolvieren bzw. absolviert haben. An dieser Stelle möchte ich auch meinen Geschwistern danken: Martin hat mich durch seinen erfolgreichen Abschluss der HTL inspiriert, es ihm gleich zu tun. Durch seinen vor kurzen erfolgten Doktorratsabschluss hat mir Bernhard den letzten notwendigen Motivationsschub geleistet. Silvia hat die Rechtschreibung der Dissertation maßgeblich verbessert und mir auch öfters durch die Betreuung meiner beiden Söhne Valentin und Jonas einen Freiraum zum Verfassen dieser Arbeit geschaffen. Auch meine Lebensgefährtin Ulli hat mich auf vielfältige Weise unterstützt. Trotz sehr turbulenter Zeiten, bedingt durch Hausbau und die Geburt unserer energiegeladenen Söhne, hat es mir Ulli immer wieder ermöglicht ausgeschlafen an der Dissertation zu arbeiten. Dafür und für ihren Rückhalt bin ich ihr äußerst dankbar.

Auch meinem beruflichen Umfeld bin ich zu großem Dank verpflichtet. Über die Jahre haben sich viele Kollegen, darunter Thomas Berer, Hubert Grün, Peter Burgholzer, Armin Hochreiner, Karoline Felbermayer, Gregor Langer, Heinz Roitner, Klaus-Dieter Bouchal, Elisabeth Leiss-Holzinger, Istvan Veres und Astrid Höllinger, bei der RECENDT mit der photoakustischen Bildgebung beschäftigt. Sie haben Grundlagen für meine Dissertation geschaffen und mit Ratschlägen, Hinweisen, Diskussionen, erfolgreichen Forschungsanträgen, Messungen, Software, Phantombau und Geräteherstellung zum Gelingen dieser Arbeit beigetragen. Mein besonderer Dank gilt Thomas Berer, der maßgebliche Betreuungsarbeit für meine Forschung geleistet hat. Auch meinen akademischen Partnern von der Universität Graz, darunter Univ.-Prof. Dr. Günther Paltauf, Dr. Robert Nuster, Dr. Klaus Passler und Dr. Sibylle Gratt, und Univ.-Prof. Dr. Markus Haltmeier von der Universität Innsbruck bin ich für die gute Zusammenarbeit dankbar. Schließlich möchte ich mich noch bei Univ.-Prof. Dr. Bernhard Jakoby für die hervorragende Betreuung, konstruktiven Diskussionen und die Möglichkeit als externer Doktorrand auf seinem Institut zu studieren bedanken.

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Zusammenfassung

Die photoakustische Bildgebung (PAB) ist ein aufstrebendes Verfahren für biomedizinische und medizinische Anwendungen. In der PAB werden Bilder von lebendem Gewebe und biologischen Proben aus Ultraschallwellen rekonstruiert. Die Schallsignale werden dabei durch optische Absorption innerhalb der Probe erzeugt und außerhalb detektiert. Unter PAB versteht man also Hybridverfahren, die vom vielfältigen optischen Kontrast und der relativ geringen Streuung von Ultraschall in Biomaterialien profitieren. Verschiedene Arten von Gewebeeigenschaften (z. B. anatomische, metabolische, histologische, funktionelle Eigenschaften) zeigen einen endogenen optischen Absorptionskontrast und können daher mit PAB untersucht werden. Durch die Verwendung von Kontrastmitteln können auch molekulare und zelluläre Informationen gewonnen werden. Ein breites Größenspektrum, das von der Zellen- bis zur Organebene reicht, kann durch verschiedene Implementierungen von PAB abgedeckt werden. Bei der PA-Computertomographie (PACT) wird eine Anordnung einer Vielzahl von Ultraschallwandlern verwendet, um rasch die notwendigen Ultraschallinformationen zu erfassen, um ein 3D-Bild eines Volumens rekonstruieren zu können. Bei der PA-Mikroskopie (PAM) wird ein 2D- oder 3D-Bild durch Rasterabtastung einer Probe mit fokussierter Ultraschallerzeugung und -detektion gewonnen.

Optische Methoden zur Ultraschalldetektion sind eine vielversprechende Alternative zu den üblicherweise verwendeten piezoelektrischen Wandlern. Sie bietet z.B. einen breiten Akzeptanzwinkel, eine hohe Empfindlichkeit, geringes Übersprechen und berührungslose Detektion. Ultraschallsensoren basierend auf faseroptischen Mach-Zehnder-Interferometern wurden im Zuge dieser Dissertation entwickelt und untersucht. In einem derartigen Detektor befindet sich ein Arm des Interferometers (d.h. die Messfaser) im zu messenden Ultraschallfeld. Der Druck entlang der Messfaser beeinflusst aufgrund des elasto-optischen Effekts ihre optische Weglänge. Werden andere Einflüsse auf die optische Weglänge (z. B. durch thermische Veränderungen) unterdrückt, variiert die Ausgangsleistung des Interferometers gemäß dem integrierten Ultraschalldruck entlang der Messfaser. Diese sogenannten integrierenden Liniendetektoren können aufgrund der mechanischen Flexibilität der Fasern in unterschiedliche geometrische Formen gebracht werden. Abhängig von der jeweiligen Form eignen sie sich besonders für spezielle Implementierungen der PAB. Zum Beispiel ermöglicht die von integrierenden ringförmigen Liniensensoren bereitgestellte hohe Tiefenschärfe photoakustische Raster-Makroskopie (PARM). In PARM werden Bilder mit einer Rasterabtastung wie in PAM erzeugt. Jedoch werden vergleichbare Abbildungstiefen und Auflösungen wie mit PACT erreicht. Für diese Dissertation wurde ein Demonstrator für PARM mit ringförmigen faseroptischen Detektoren entwickelt und charakterisiert. Die laterale Auflösung wurde auf 150 µm bis 200 µm geschätzt, basierend auf der Vermessung eines Blattskeletts unter einer 10 mm dicken Schicht aus Agarose. Die laterale Auflösung bei einem axialen Abstand von 110 mm erhöht sich nur um den Faktor 1,8 gegenüber der Auflösung bei 15 mm Achsabstand.

Eine weitere für diese Dissertation entwickelte PAB-Implementierung verwendet 64 faseroptische parallel verlaufende gerade Liniendetektoren, die entlang eines Kreises angeordnet sind. Sie ermöglicht die photoakustische Projektionsabbildung (PAPA) von Objekten mit einer Größe von mehreren Zentimetern. Eine Bildauflösung im Bereich von 100 µm bis 260 µm wurde erreicht. Das

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3D Problem der Beobachtung größerer Volumina kann mit PAPA auf ein 2D Problem zurückgeführt werden. Das bringt den erheblichen Vorteil einer Reduktion der notwendigen Sensorpositionen um zwei Größenordnungen.

Die entwickelten Demonstratoren für PAPA und PARM erzielen ansprechende Bildauflösungen und Empfindlichkeiten. Das Auftreten von Abbildungsartefakten sollte jedoch durch eine Erhöhung der Anzahl von Sensoren reduziert werden. Des Weiteren sollte die Abbildungsgeschwindigkeit durch die Verwendung eines schnelleren Anregungslasers, durch paralleles Auslesen aller Messkanäle und durch Änderung des Scanvorgangs beim PARM-System erhöht werden. PAB-Systeme, die auf faseroptisch integrierenden Liniendetektoren basieren, können eine vielversprechende Alternative und Erweiterung zu bestehenden PAB-Verfahren sein.

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Abstract

Photoacoustic imaging (PAI) is an emerging method for biomedical and medical applications. In PAI, images of living tissue and biological samples are reconstructed from ultrasonic waves generated by optical absorption and detected on the outside. Thus, PAI denotes hybrid methods benefiting from the diverse optical contrast and the relative low scattering of ultrasound in biomaterials. Different types of tissue properties (e.g. anatomical, metabolic, histological, and functional properties) show endogenous optical absorption contrast and can therefore be examined with PAI. By using contrast agents, also molecular and cellular information can be gained. Object sizes ranging from cellular to organ level are covered by different implementations of PAI. In photoacoustic (PA) computed tomography (PACT), an array of ultrasound transducers is used to quickly acquire the necessary ultrasound information to reconstruct a 3D image of a volume. In PA microscopy (PAM), a 2D or 3D image is obtained by raster scanning a sample with focused ultrasound generation and detection.

Optical methods are a promising alternative to commonly used piezoelectric transducers (PET). They offer e.g. a wide acceptance angle, high sensitivity, low-cross talk, and remote detection. Ultrasound sensors based on fiber-optic Mach-Zehnder interferometers were developed and investigated in the course of this thesis. In such a detector, one arm of the interferometer (i.e. the measurement fiber) is exposed to the ultrasonic field. The pressure along the measurement fiber affects its optical path length due to the elasto-optic effect. By suppressing other influences on the optical path length (e.g. thermal drifts), the output power of the interferometer varies according to the integrated ultrasonic pressure along the measurement fiber. These so-called integrating line detectors can be brought into different geometric shapes due to the mechanical flexibility of the fibers. Depending on the respective shape, these sensors allow special implementations of PAI. For example, the large depth-of-field offered by annular integrating line detectors makes PA scanning macroscopy (PASMac) feasible. In PASMac images are acquired with a scanning approach like in PAM but with imaging depths and resolutions comparable to PACT. For this thesis, a demonstrator for PASMac with annular fiber-optic detectors was developed and characterized. The lateral resolution was estimated to be 150 µm to 200 µm based on the imaging of a leaf skeleton under a 10 mm thick layer of agarose. The lateral resolution at an axial distance of 110 mm increases only by a factor of 1.8 compared to the resolution at 15 mm axial distance. Another PAI implementation developed for this thesis features 64 fiber-optic parallel straight line detectors arranged along a circle. It facilitates PA projection imaging (PAPI) of objects with several centimeters in size. An imaging resolution in the range of 100 µm to 260 µm was achieved. The 3D problem of monitoring of large volumes can be reduced two a 2D problem by PAPI. This brings the major advantage of reducing the number of necessary sensor positions by two orders of magnitude.

The demonstrators developed for PAPI and PASMac achieve respectable imaging resolutions and sensitivities. The occurrence of imaging artifacts should be reduced by increasing the number of sensors. Furthermore, the imaging speed should be increased by using a faster laser, by parallel readout of all measurement channels and by a modification of the scanning process in the PASMac system. In conclusion, PAI systems based on fiber-optic integrating line detectors can be a promising alternative and extension to existing PAI methods.

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1 Motivation

Photoacoustic imaging (PAI) denotes hybrid methods using photoacoustically generated ultrasound to create images with optical contrast of biological structures (see Section 2.1.1 and [1]–[3]). Ultrasound is generated in the structures to be imaged by transient heating with intensity modulated light, most often with short laser pulses. The generated ultrasonic signals are then detected outside the object and used to form an image of the optical absorption. The imaging depth can be larger compared to pure optical methods (e.g. diffuse optical imaging, fluorescence imaging, and optical microscopy) because acoustic scattering in tissue is three orders of magnitude lower compared to optical scattering [1], [4]. PAI is well suited to be a successful complement to and replacement of established imaging modalities [5]. Thus, considerable effort is put into the development and commercialization of various types of PAI (see Section 2.1.3 and [6]). PAI devices are applied in pre-clinical studies (e.g. [7]–[9]) and in biomedical laboratories and increasingly find their way into standard medical operation (e.g. [10]–[13]). By now, some PAI devices are already commercially available [14]–[16].

Usually, piezoelectric transducers (PET) are used to detect ultrasound. In photoacoustic microscopy (PAM, see Section 2.1.3.2), a focused single element transducer acquires the ultrasonic response from a small volume. Raster scanning then yields a 2D or 3D image. In photoacoustic computed tomography (PACT, see Section 2.1.3.1), the ultrasonic field around the sample is captured by PET arrays and used to reconstruct a 3D image. PET are applied for PAI because they are easily accessible, well established and directly convert the ultrasonic signals into electric signals. Moreover, the use of transducer arrays built for ultrasonography ensures relatively low costs. However, PET are not a perfect match for PAI. The imaging resolution in PAI depends among other effects on the bandwidth of the acquired PA signals. In many cases, especially in PAM, the required bandwidth to fully detect the broadband photoacoustic (PA) signals (see Section 2.1.2 and [17]) cannot easily be achieved with PET. Also, the imaging quality, resolution and artefact suppression are negatively affected by the relatively narrow acceptance angle of PET [17], in particular in PACT. Furthermore, PET are typically opaque and thus are hindering the PA excitation. Moreover, cross-talk and electromagnetic disturbances can cause problems. Since the sensitivity of PET decreases with the size of the sensing element, miniaturization of PET arrays is limited. In contrast, piezoelectric micromachined transducers (PMUTs [18]) and capacitive micromachined transducers (CMUTs [19]–[21]) can be miniaturized and show enhanced bandwidths.

As another alternative to PET, optical ultrasound detection [17], [22], [23] could meet the requirements of PAI better. Usually, those techniques can easily provide a proper detection-bandwidth. In addition, optical methods can provide high sensitivity (i.e. a high signal-to-noise

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unhindered sample illumination and the possibility of miniaturization. However, since optical ultrasound detectors are much less established compared to PET, the cost-effectiveness must be addressed when a commercial product is aimed for [26].

For some PAI applications, e.g. PA projection imaging (PAPI, see section 2.2.1 and 3.2) and PA macroscopy (PASMac, see Section 2.2.2 and 3.3), detectors which integrate the pressure along one dimension are particularly well suited. Such transducers may be realized e.g. with PET [27], free beam interferometry [28] or fiber-optic interferometers [26], [29]–[31]. The PET approach is subject to limitations of PET detectors already mentioned above and the free beam technique cannot be straightforwardly parallelized to yield an array of detectors and this approach also exhibits an undesired varying diameter of the sensing element (i.e. the laser beam). Fiber-optic interferometers can overcome these limitations and match the requirements of PAPI and PASMac. Such detectors offer a low noise-equivalent pressure over a wide frequency range without showing directionality of the sensitivity. Since the diameter of the sensing element (i.e. the guided laser beam) is maintained by the fiber, the bandwidth of the detector is constant along its length. By using fiber-optic equipment available from telecommunication industries, the fibers can be used to yield relatively low-cost sensor arrays [26]. Additionally, the transparent fibers do not block the PA excitation. Moreover, since the interrogating laser beams are guided by fibers, the read-out (e.g. by balanced photodetectors) of the interferometers can be spatially separated from the sample. Hence, miniaturization of the sensor is not hindered by the read-out system. Because the fibers are flexible, they can be shaped into different geometries. Thereby, the path along which the fiber-optic detectors integrate the PA pressure can be matched to the desired application (e.g. line detectors for PAPI and ring detectors for PASMac).

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2 Introduction

2.1 Photoacoustic Imaging

2.1.1 Principle

A photoacoustic (PA) image represents an image of volumetric energy density generated by absorbed photons. Figure 1 shows the principle of PAI. First, absorbers in the sample are heated by intensity modulated electromagnetic waves. The absorbers emit acoustic waves if certain conditions are fulfilled (see Section 2.1.2). These waves are detected after leaving the sample. Finally, the initial pressure distribution (representing the image of the energy density) is reconstructed on basis of the measured sound waves. The main implementations of PAI are presented in Section 2.1.3.

2.1.2 Generation and Propagation of Photoacoustic Signals

PAI makes use of the PA or optoacoustic effect [32], [33] which describes the generation of sound waves caused by absorption of light. This effect is not limited to the visible spectrum of electromagnetic radiation. However, most implementations of PAI use visible to near- infrared light because biologically meaningful substances (e.g. hemoglobin, lipid, water, bilirubin, melanin, cytochrome c, and myoglobin [34]) absorb within this frequency range. Moreover, such radiation is non-ionizing and thus carries less harm to health.

The PA generation of sound waves can be described by the following sequence [35]: First, electromagnetic radiation is absorbed within a sample depending on the local absorption coefficient. Then, most of the absorbed energy is converted into heat. This leads to a temperature rise and subsequent thermoelastic expansion. When the electromagnetic radiation is not constant (e.g. by using pulsed laser or intensity modulated laser illumination), this thermoelastic expansion

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< _{= 4 ,} (1) denotes the size of the heated volume and represents the thermal diffusivity and in

< = , (2)

is the speed of sound. In tissue, is usually much smaller than . Structures smaller than ∙ generate weak signals and will appear blurred in the resulting image, since the generated ultrasound can escape from the heated volume during the heating pulse. When designing a PAI system, should be set to ⁄ , where denotes the size of the smallest absorber, which should be resolved by the system. The fractional volume expansion of the heated sample at position =

is given by

= − + ! , (3)

where is the pressure rise, denotes the isothermal compressibility, represents the thermal coefficient of volume expansion and ! is the increase in temperature. If thermal and stress confinement are fulfilled, the fractional volume expansion / is negligible during heating [36] and eq. (3) can be used to calculate the initial pressure rise

$ = ! . (4)

For tissue, / is approximately 8 ⋅ 10*+, -⁄ [37]. Hence, already a temperature rise of a millikelvin generate practical ultrasound signals. If nonthermal effects, e.g. fluorescence, can be neglected, the initial pressure rise can be expressed by

$ = _./

01 , (5)

where 1 represents the energy absorbed per volume, /₀ denotes the specific heat capacity and . is the mass density. The factor / ./₀ is called Grüneisen parameter. The resulting ultrasonic field in a homogenous lossless medium can be described by the partial differential equation (PDE)

2∇ − 1 4_{45 6} , 5 = 0, (6)

where ∇ represents the Laplace operator 4 47⁄ + 4 48⁄ + 4 49⁄ . The initial conditions of this PDE are given by

, 0 = $ and ₄₅4 , 0 = 0 (7)

Wang shows in [37] how the PDE (6) with initial conditions (7) can be solved e.g. by Green’s function approach. The resulting PA signal of a spherical absorber which is homogeneously heated by a delta pulse is given by

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Here, the center of the sphere is located at the coordinate origin. By using the radial symmetry of the problem, the PA pressure , 5 is a function of the distance to the origin and time 5. The initial pressure distribution _$ is given by

$ = ̌$< < =>− , (9)

where < represents the Heaviside step function, which is 0 for < 0 and 1 everywhere else. Hence, the initial pressure has a value of ̌_$ within the radius of the sphere =_> and 0 outside the sphere. Analyzing eq. (8) reveals the following insights:

The initial pressure distribution leads to the emission of two waves with a positive pressure amplitude: (i) a spherical wave propagating in negative radial direction and (ii) a spherical wave travelling in radial direction.

The amplitude of the inwardly launched wave (ii) changes its sign after passing the center of the sphere and continues propagation in radial direction.

The amplitude of the waves outside the sphere is inversely proportional to the distance from the sphere.

The resulting N-shaped PA signals are shown at three distances to the origin in Figure 2. The duration of the pressure pulse equals the time it takes for a sound wave to pass through the sphere. Thus, the bandwidth of PA signals is inversely proportional to the absorber size. Sources with a diameter within the range of 10 µm generate signals with a bandwidth above 150 MHz. Actual PA signals generated e.g. in tissue differ, of course, from this idealistic response. Significant effects stem from inhomogeneous illumination and acoustic attenuation. Because acoustic attenuation increases with frequency (see Section 3.1 and [38]), the bandwidth of ultrasonic signals decreases with the imaging depth (i.e. the propagation distance to the detector). For PA simulations in tissue-realistic media, the k-wave toolbox can be used [39].

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2.1.3 Main Implementations of PAI

By now, many different implementations of PAI have been developed. Often, they are designed for a specific application, but some also seek to offer a more flexible usage. The various types can be compared in terms of certain characteristics, e.g. imaging depth, resolution, field-of-view, aimed application, frame-rate and transducer type. Since a comprehensive presentation of the state-of-the-art PAI systems is beyond the scope of this thesis, the following two subsections only give a short overview of the main PAI types to which most of the PAI systems can be attributed to. More information on this topic can be found in the literature [1], [2], [40], [41].

2.1.3.1 Photoacoustic Computed Tomography (PACT)

The basic principle of PACT is shown by Figure 3. The excitation laser beam is expanded to fully illuminate the region of interest. The generated PA signals propagate to the surface of the sample and are detected by a transducer array. In a second step, the initial pressure distribution is reconstructed from the acquired data [39], [42], [43]. Thus, in principle, a single laser shot is sufficient to produce an image. However, to yield high-resolution images with few imaging artifacts, commonly used transducer arrays must perform multiple measurements at different positions to acquire sufficiently dense PA data. Also, since PA signals are generally weak, signal averaging is often necessary.

Several geometries of the transducer arrays are applied. Two dimensional arrays with e.g. planar [7], [44] or (hemi-)spherical [45]–[48] geometry inherently deliver 3D data. The number of detection elements of such arrays is technically limited to a few hundred. For adequate 3D imaging, however, the number of necessary detection positions is in the order of 104_{. Therefore, multiple} measurements with such arrays need to be performed. Linear [16] or circular arrangements [49], [50] yield cross-sectional images. By translating these 2D arrays perpendicular to the imaging plane, neighboring cross-sectional images can be acquired and stacked to form a 3D image. Also, multiple measurements around a sample with an arc-shaped array [14], [51] deliver PA data for a 3D reconstruction.

Despite the difficulties to achieve a high frame-rate in PACT, also real-time systems have been proposed [15], [49].

Figure 3: Principle of photoacoustic computed tomography (PACT). The whole region of interest of the object is illuminated. The generated photoacoustic signals are detected by a transducer array and used to reconstruct the image of the absorbed photons.

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2.1.3.2 Photoacoustic Microscopy (PAM)

Figure 4 presents the basic functionality of PAM. In contrast to PACT, only a small volume of the sample is excited by the laser beam during a single measurement. A focused ultrasound transducer is aligned to the excited spot within the sample. Thus, a single measurement yields a PA signal stemming from a small region shaped by the focusing of the laser beam and the ultrasound transducer. Typically, such a region has an elongated form. The longitudinal axis of this focus region corresponds to the z-axis in Figure 4. Hence, the lateral position in the xy-plane of the PA signals is given by the xy-position of the laser beam and the ultrasound transducer respectively. The z-position (i.e. the depth) of a PA source can be determined by multiplying the time-of-flight of the signal by the speed of sound in the sample. Thus, a single measurement in PAM yields depth-resolved optical absorption at a single lateral position. By a linear lateral movement of the focal region (by scanning the laser beam and/or the ultrasound transducer), a 2D PA image can be acquired. Several line scans can then be composed to form a 3D image. PAM can be divided in two main branches [41], [52]: In optical resolution PAM (OR-PAM) [53], [54] the focal region is mainly determined by the optical focusing of the excitation laser beam whereas in acoustic resolution PAM (AR-PAM) [55] the main contribution to the focusing stems from the ultrasound transducer. Due to optical scattering, the imaging depth in OR-PAM is usually limited to approximately 1 mm. Since acoustic scattering is much weaker than optical scattering, AR-PAM can achieve larger imaging depths. Like in PACT, the frequency-dependent acoustic attenuation and decreasing light delivery in tissue limit the practical imaging depth in AR-PAM. To enhance light delivery in deep tissue, dark field illumination is applied in AR-PAM [56].

Figure 4: Principle of photoacoustic microscopy (PAM). The object is illuminated with a focused laser beam. A focused ultrasound transducer detects the PA signal. Illumination and detection is coaxially aligned. The region of interest is raster-scanned to form an image.

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2.2 Photoacoustic Imaging with Integrating Detectors

If the size of the sensing element of an ultrasound transducer is small relative to the wavelength of an incident ultrasonic wave, the sensor is denoted as point detector. When the integration of the ultrasonic field over the sensor area significantly influences the sensor output, the transducer is called integrating detector. For instance, focusing piezoelectric transducers (PET) with a spherically shaped sensing element are most sensitive to signals stemming from the center of the sphere. However, the single elements of a PET array are considered to be rather point-like. As already mentioned in Section 2.1.3, focused (i.e. integrating) transducers are usually used for PAM and arrays of point-like detectors are most often applied in PACT. Alternatively, integrating planar [57] or line [58] receivers can be employed in PACT [59]. Special image reconstruction methods are necessary for PACT using such integrating detectors. To function properly, these methods require the integrating sensors to be several times larger than the imaged object [60]. The following two subsections introduce PAI modalities based on integrating line detectors.

2.2.1 Photoacoustic Projection Imaging with Integrating Straight Line Detectors

Photoacoustic projection imaging (PAPI) can be realized with an array of straight line detectors. The response of an individual line sensor to the N-shaped pulse of a point-like PA source (see Section 2.1.2) can be described in three phases. Figure 5 schematically shows the occurring pressure along the sensor in these phases. At first, the ultrasonic pulse only partly overlaps with the line detector in the region around the normal distance to the source, marked by (1) in the figure. Since the sensor integrates the pressure along its length, the resulting signal in this phase initially shows a positive peak. The integrated signal begins to decrease when the negative part of N-shaped pulse reaches the sensor. When the wavefront propagates further, it intersects in two zones with the line sensor in phase two. In these regions, marked by (2) in the figure, the pressure exhibits a complete N-shaped form. Due to the form of the pulse, the positive part of the signal overlaps with the line sensor at a larger distance to the source than the negative part. Because the amplitude of the pressure wave decreases with distance to the source, the negative amplitude is larger than the positive amplitude. Hence, the response of the integrating line sensor is negative in phase two. For distances several times larger than the diameter of the source, the difference of positive and negative amplitude of the pulse becomes very small. When the wavefront crosses the end of the sensor in phase three, the positive part of the N-shaped pressure pulse begins to leave the sensor. Thus, the integrated signal increasingly becomes negative. With increasing length of the line sensor, this negative peak will appear later and decrease in amplitude.

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Figure 6 shows the measurement configurations for PAPI. The sample is placed in the middle of the array of circularly distributed parallel line sensors. The whole sample is excited with short laser pulses (not shown in the figure). Thus, all PA sources in the sample simultaneously emit PA signals, which are received by the sensor array. To avoid imaging artifacts, the negative sensor response occurring when a PA pulse reaches the end of the sensor (see phase three in Figure 5) must not overlap with any initial response (corresponding to phase one in Figure 5). This is the case when the length of the line detectors is larger than 2 + ?, where denotes the diameter of the array and ? the sample length in z-direction. Then, a line sensor will be reached by a pulse from a PA source with the maximal possible normal distance (i.e. from a distance D) before a pulse from a PA source close to the sensor will reach the end of the line sensor. Note that the sample has to be placed in the middle of the array. Hence, all possible unwanted signal contributions from pulses reaching the end of the array arrive at 5 @ _>/ and can be cut off without losing any wanted signals from the initial sensor responses [60]. Such processed signals are

Figure 5: Propagation of an N-shaped PA pulse stemming from a point-like source (red). The passing of the wave through the line sensor (green) is shown in three phases (numbers 1-3 in the cycles). The pressure along the sensor at three different times t1 to t3 (corresponding to the three phases) is presented in the graph (c) above the line sensor.

The time traces of the pressure are shown for each phase on two different positions on the sensor in (a). The resulting integrated pressure is presented in (b).

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to a 2D problem. This brings the major advantage of reducing the number of necessary sensor positions by two orders of magnitude. Like in X-ray computed tomography, a 3D image of the sample can also be reconstructed from several projections from different angles.

2.2.2 Photoacoustic Scanning Macroscopy with Integrating Annular Detectors

Figure 7 presents a point-like PA source relative to an annularly shaped integrating detector. If the source is on the ring axis, the PA signal reaches the entire sensor simultaneously. Thus, sources on the ring axes generate significantly stronger signals in ring-shaped sensors compared to straight line detectors with equal sensor length and distance to the source. The situation for sources lying off the axis is comparable to phase one and two (see Section 2.2.1) of the response of a straight line detector to an N-shaped PA pulse. A significant sensor response to off-axis sources is only generated when the wavefront initially reaches the ring sensor after propagating AB and when the wavefront leaves the sensor after propagating AC. During these phases, the

N-shaped PA pulse is only partly present along the ring. The generated signals in these phases are, however, much weaker than signals from roughly equally strong PA sources lying on the ring axis because a much smaller part of the ring sensor interacts with the PA pulse. When the whole N-shaped PA pulse is present along the ring (i.e. when the wavefront crosses the ring in two zones), the integrated pressure is virtually zero. Hence, an annularly shaped integrating detector is, by far, most sensitive to signals stemming from the ring axis but also responds to off-axis sources. Because the focusing mechanism of integrating ring transducers is not affected by the axial distance of the source, such detectors offer a very large depth-of-field compared to focusing spherical sensors.

Figure 6: Arrangement of integrating line detectors for photoacoustic projection imaging (PAPI). The line-shaped sensing elements (colored in green) are aligned along the z-axis and are uniformly distributed along a cycle with diameter D. By using adequate PA reconstruction methods [60], a projection of the PA absorbers (red) within the sample (gray) onto the xy-plane can be generated from the measured PA signals.

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The axial distance of an absorber on the ring axis can be calculated by D _>5 − = , where 5 is the time-of-flight of the signal and = represents the ring radius (see Figure 7a). With this transformation, the ring-shaped transducers can be used for PA microscopy as explained in Section 2.1.3.2. The off-axis sensitivity of the annular sensors can cause imaging artifacts. When a ring sensor with radius =_E is laterally scanned over a single absorber lying at a depth of , two peaks appear in the measured signal as long as the absorber lies off the ring axis. Figure 8a illustrates this situation. One peak will arrive earlier after propagating A_B= D =_E− ,_F − and the other peak will arrive later after propagation A_C= D =_E+ ,_F − . Here, ,_F represents the normal distance from the PA source to the ring axis. Figure 8c shows the signals as function of the pulse propagation distance, A = _>5, for two concentric ring sensors with different diameters for varying ,_F. The projection of these signals on the ring axis (see Figure 8d) results in artefact PA sources at a depths of

B,G= HAB− =G and C,G= HAC− =G

with i ∈ K1,2L. The smaller the distance of the PA source from the ring axis is, the less the two peaks are separated in the signal response of a ring. If the PA absorber lies on the ring axis, the two peaks merge to one signal with several times higher amplitude. Note that the depth at which the artefact peaks appear depends on the ring radius. Thus, the artifacts can be significantly reduced by superposing the signals of multiple concentric ring detectors with varying ring radius. This artefact suppression is usually enhanced by coherence weighting [61]. Such an array of integrating ring sensors is applied in our system for PA scanning macroscopy (PASMac). This

Figure 7: Annularly shaped integrating detectors (green) with PA point-like source (red) on (a) and off (b) the ring axis. The intersection of the wavefront with the ring plane is shown by the dotted red cycles at different times. In (a) the wavefront reaches the entire ring simultaneously whereas in (b) only parts of the ring interact with the wavefront at a time. The sensor only generates a signal from off-axis PA sources when the wavefront initially reaches the sensor after propagating s- and when the wavefront leaves the sensor after propagating s+.

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2.3 Fiber-Optic Realization of Integrating Line Detectors for PAI

The integrating line detectors introduced in the previous section can be implemented on basis of fiber-optic Mach-Zehnder interferometers (MZI). The PAI prototypes I developed for this thesis (see Section 3.2 and 3.3) use such detectors. Section 2.3.3 describes their design, composition and functionality. The operation of a MZI and its fiber-optic implementation are presented in Section 2.3.1. The interaction of ultrasonic wave and the interferometer is discussed in Section 2.3.2.

2.3.1 Mach-Zehnder Interferometer

The schematic illustration of a typical MZI is shown in Figure 9. A continuous wave laser beam is split into a reference and measurement beam by a beam splitter. Each beam is directed by a mirror onto a second beam splitter. Two beams, each consisting of a superposition of reference and measurement beam, exit the second beam splitter. If the interferometer is properly operated, the difference of the optical path of reference and measurement beam between the beam splitters is below the coherence length of the laser. Thus, interference occurs in the output beams. The phase difference of measurement and reference beams caused by the mirrors and beam splitters is M at one output and 2M at the other output. An additional phase shift in the measurement beam of ΔO alters the intensity of the output beams according to

+E~QRA ΔO/2 and + ~AST ΔO/2 . (10)

Here, +_E and + denote the intensities at output one and two respectively. The additional phase shift ∆O results from a variation of the refractive index along the path and/or different path length. Equations (10) can be derived from the fact that the intensity is proportional to squared amplitude of the electric field. For example for output one, the amplitudes of the electric field of reference and measurement beam have to be vectorially added.

Figure 8: (a) An integrating ring sensor with radius R1 responds to a pulse stemming from a PA source lying ao off

the ring axis when the wavefront of the pulse initially reaches the ring after propagating s_ and when the wavefront

leaves the sensor after travelling s+. The projection of these sensor responses on the ring axis results in two artefact

PA sources lying at d- and d+. (b) shows the depths of the artifact sources when two concentric rings with diameters

R1 and R2 are used to detect the PA signals. (c) presents the signals of the two ring sensors as function of the pulse

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Hence,

+E~VWXYV = VWXXXXXXY + WZE XXXXXXXYV [E

where WXXXXXXY and W_ZE XXXXXXXY represent the amplitude of the electric field at output one of reference and _[E measurement beam respectively. If the coherence length of the laser beam is larger than \ 1 + ΔO/2M , which is usually the case, the above statement can be expressed by

+E~V]^_+ ]^`_C^abV = V1 + ]^abV = c1 + QRA ΔO d + AST ΔO

1 + QRA ΔO + AST ΔO + 2QRA ΔO = 2c1 + QRA ΔO d = 2QRA ΔO/2

Figure 10 shows the principal schematic of the fiber-optic implementation of the MZI. Here, fiber couplers take over the task of the beam splitters. These couplers mostly work on basis of resonant coupling [62], where, in the simplest case, two fibers are brought into close proximity to each other. Then, a beam in one fiber partly couples into the other fiber. Depending on the geometry of this coupling zone, the power transfer rate can be adjusted relatively freely. To achieve the most sensitive MZI, the intensities of the superimposing light beams should be equal. If the losses of light are equal in the reference and measurement arm, the intensities of the superimposing beams can be matched by choosing an equal coupling ratio for both fiber couplers. In the coupled beam (i.e. the beam which is transferred from one fiber into the other), a phase shift of M/2 occurs. In the beam that continues in the original fiber (i.e. the input fiber), no phase shift occurs. The complex amplitudes of the electric field at several positions are noted in the Figure 10. The relative phase shift between the paths before the second fiber coupler without the contribution of the first coupler is denoted ∆O. Thus, ∆O results from differences in the optical path lengths between the couplers. The intensity of the output beams is, similar to the free-beam MZI, proportional to QRA ΔO/2 and AST ΔO/2 respectively.

Figure 9: Schematic representation of a Mach-Zehnder interferometer realized with mirrors and beam splitters. The phase shifts caused by the mirrors and beam splitters are noted next to the respective part of the beams and are indicated by the color of the beam. In the measurement arm, an additional phase shift ∆O occurs. Except for this phase shift, the optical path length of reference and measurement beam is a multiple of the wavelength. The resulting beam intensity depends on ∆O . The proportionality is given next to the output beams.

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2.3.2 Interaction of an Acoustic Wave with an Optical Fiber

An acoustic wave passing through an optical fiber induces strains in the fiber. According to the elasto-optic effect [63], these strains alter the refractive index (RI) in the fiber. If the wavelength of the acoustic wave is large compared to mode-field diameter of the guided beam, the acoustic pressure at a certain longitudinal position A can be approximated to be constant over the beam cross-section. Thus,

T A = T$ A + Tc A d, (11)

where T_$ A denotes the RI in absence of acoustic pressure and Tc A d represents the variation of the RI due to the acoustic pressure at location A. If the fiber is birefringent, the RI may be different for the two polarization components. For a basic understanding of the interaction of an acoustic wave with an optical fiber, the influence of the polarization can be disregarded. The optical path length Λ of an optical fiber running along the curve / is given by

Λ = f T A A.

g

(12) Thus, a measure of the integrated acoustic pressure along the fiber can be given by

f Tc A d A = Λ − f T$ A A = Λ − Λ$. g

g

(13) A fiber-optic integrating line detector for acoustic pressure can then be realized by measuring the phase shift caused by the variation of the optical path length. Berer et al. presented a sophisticated description of the response of fiber-optic integrating line detectors to ultrasonic waves [64]. This work includes theoretical and measured transfer functions depending on incident angles and ultrasonic frequency for glass and polymer fibers. An important finding was that polymer fibers exhibit a significantly stronger response to phase shifts and a more uniform frequency behavior. Also, the response of polymer fibers is virtually independent from the polarization state of the guided beam. On basis of these results, polymer fibers were chosen as pressure sensing element in the detectors developed for this thesis.

Figure 10: Schematic representation of a fiber-optic Mach-Zehnder interferometer. A fiber coupler branches the source beam in a measurement and a reference path. A second fiber coupler superimposes the light in these paths into two output fibers. The complex amplitudes of the electric field of the guided beams at different locations are stated. Like for free-beam MZI, the intensities of the output beams depend on the phase shift between measurement and reference path.

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2.3.3 Design of an Ultrasound Detection System based on Fiber-Optic Mach-Zehnder Interferometers

Figure 11 presents the schematic composition of an integrating line detector for PAI based on a fiber-optic Mach-Zehnder interferometer (FOMZI) explained in Section 2.3.1. To get a functional sensor, the MZI needs a read-out system and a stable working-point phase shift (WPPS). The read-out system consists of a balanced photodetector (BPD), an analog-digital converter (ADC) and a personal computer (PC). The BPD has a high-frequency (HF) output, carrying only the PA signals. The ADC samples the PA signals and feeds them to a PC for recording. The stable WPPS is established by a phase shifter in the reference path. Disturbances of the WPPS with frequencies below the frequency range of the PA signals, e.g. caused by thermal drifts, are available at the low-frequency (LF) output of the BPD. By feeding the LF-signal to a controller, driving the phase shifter, such disturbances can be suppressed.

To achieve best sensitivity, the WPPS between reference and measurement arm should be M/2. Then, the beams at both outlets of the MZI have equal intensity, and the BPD, converting the intensity difference at its inputs to an electric voltage, generates zero signal. Thus, the set-point of the controller is zero volt. Starting from this WPPS, a phase shift in the measurement path, induced by a PA signal passing through the polymer fiber (see previous section), affects thereby both outputs of the MZI equally in magnitude. The difference from the WPPS is, however, positive at one output and negative at the other. Figure 12 shows the intensity at the outlets of the MZI as function of the phase shift. Also, the response to an N-shaped PA signal is included.

Figure 11: Schematic functionality of an integrating line detector for PAI based on a fiber-optic Mach-Zehnder interferometer. The first fiber coupler equally splits the output of a continuous wave (cw) laser into the reference and measurement arm of the Mach-Zehnder interferometer. A phase shifter establishes a working-point phase difference of measurement and reference arm of M/2. A second fiber coupler superimposes the beams propagating in the reference and measurement path. A balanced photodetector converts the light into high (HF) and low frequency (LF) electric signals. The HF signal carrying the photoacoustic signals (red) is sampled by an analog-digital converter and subsequently recorded by a PC. The LF signal contains disturbances of the working-point, e.g. by thermal drift, and is fed to a controller which drives the phase shifter.

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Figure 12: Intensity at the outputs of a FOMZI as function of the phase shift between the light beams propagating in the reference and measurement path. When the FOMZI operates at a working-point phase shift of M/2, the response to PA induced phase shifts is optimal.

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3 Published Articles

This section contains three peer-reviewed articles published in scientific journals. The essential works and results of my dissertation are presented in these articles. I was the principal author of all three articles. Each article is presented in a separate subsection. At the beginning of each subsection, the motivation for the work is briefly explained and the contribution of the respective authors is noted.

3.1 TUFFC: Broadband High-Frequency Measurement of Ultrasonic

Attenuation of Tissues and Liquids

Since the frequency-dependent acoustic attenuation leads to a low-pass filtering of the PA signals and therefore limits the imaging depth (the imaging resolution, respectively), the expected acoustic attenuation should be taken into account when designing a PAI implementation. Also, the PA reconstruction can be enhanced if the frequency dependency of the attenuation is known [65]. Since not much information could be found in the literature, we decided to carry out our own measurements. I developed the concept of the measurements, conducted the experiments and did the data analysis. All authors provided critical feedback and contributed to the final manuscript.

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IEEE TRANSACTIONS ON ULTRASONICS, FERROELECTRICS, AND FREQUENCY CONTROL, . 59, . 12, DECEMBER 2012 2631

Broadband High-Frequency Measurement of

Ultrasonic Attenuation of Tissues and Liquids

Johannes Bauer-Marschallinger, Thomas Berer, Hubert Grün, Heinz Roitner, Bernhard Reitinger, and Peter Burgholzer

Abstract—The ongoing expansion of the frequency range used for ultrasonic imaging requires increasing attention to the acoustic attenuation of biomaterials. This work presents a novel method for measuring the attenuation of tissue and liq-uids in vitro on the basis of single transmission measurements. Ultrasound was generated by short laser pulses directed onto a silicon wafer. In addition, unfocused piezoelectric transducers with a center frequency of 50 MHz were used to detect and emit ultrasound. The laser ultrasound method produces signals with a peak frequency of 30 MHz. In comparison to piezoelec-tric generation, pulse laser excitation provides approximately 4 times higher amplitudes and 20% larger bandwidth. By using two excitation methods in succession, the attenuation param-eters of porcine fat samples with thicknesses in the range of 1.5 to 20 mm could be determined quantitatively within a total frequency range of 5 to 45 MHz. The setup for liquid measure-ments was tested on samples of human blood and olive oil. Our results are in good agreement with reports in literature.

I. I

M

 for measuring ultrasonic attenuation in

bio-logical tissues and liquids have been a subject of scientiﬁc research since the development of medical ultra-sonic imaging. These measurements are normally based on piezoelectric transducers with a narrowband frequency spectrum of up to 10 MHz. Attenuation data for higher frequencies are rare; however, they are of interest for re-cent developments in the ﬁeld of medical ultrasonic imag-ing. This work presents a method for measuring acoustic attenuation up to frequencies of 75 MHz on the basis of ultrasonic generation by short laser pulses and piezoelec-tric transducers.

The information loss resulting from acoustic attenua-tion determines how far inside the body images can be acquired and what resolution those images may achieve. Images of tissue from deep within the body may have less resolution than images from beneath the skin, because ultrasonic attenuation in the subcutaneous fat layer

in-creases with frequency [1]. Because high-frequency appli-cations in ultrasonic imaging are currently under develop-ment, strength and frequency dependency of attenuation beyond 10 MHz are increasingly of interest. For example, traditional pulse–echo ultrasound has been extended to frequencies up to 100 MHz and used to investigate the human dermis [2]–[4]. A promising technology called pho-toacoustic tomography (PAT) [5], [6] also operates with broadband and high-frequency ultrasonic signals. PAT combines the advantages of diﬀuse optical imaging (high contrast) and ultrasonic imaging (high spatial resolution). In addition to many other factors, the accuracy of the acquired images depends on acoustic attenuation [7]. It is not very surprising that reconstructed images from at-tenuated signals have less contrast and lack sharpness. For example, in the case of early detection of breast cancer with PAT, small tumors in the mammary gland might not be visible without compensation of attenuation. Models of acoustic attenuation in photoacoustic imaging and pos-sible compensation methods on the basis of Szabo’s equa-tion [8] are proposed by, e.g., Roitner and Burgholzer [9] and La Rivière et al. [10]. Another approach developed by a group at the University College of London [11], [12] uti-lizes a modiﬁed wave equation with fractional Laplacian [13]. Of course, the performance of compensation depends on the noise level in the measured signals and an accurate knowledge of the attenuation parameters.

Raju and Srinivasan [3] and Bamber [14] give an over-view of studies concerned with ultrasonic attenuation of fat tissue. The attenuation parameters are well known for frequencies up to 10 MHz (see, e.g., [15], [16]). However, for higher frequencies, not quite as many investigations (e.g., [2], [17], [18]) have been conducted. This work con-tributes to the development of measuring methods of at-tenuation of tissues and liquids within the frequency range utilized by PAT and other high-frequency techniques. Ad-ditionally, it extends the frequency range for which in

vi-tro attenuation of porcine tissue has been investigated to

45 MHz.

The presented method uses laser-generated ultrasound [19], [20] and unfocused, broadband, high-frequency piezo-electric transducers. Laser ultrasound has long been used as a contactless investigation method for material test-ing and characterization up to frequencies of several hun-dred megahertz. Oraevsky et al. used laser-generated ul-trasound to measure optical properties of the tissue [21], however, there is no study known to the authors which uses laser ultrasound for measuring acoustic attenuation of tissue.

Manuscript received June 26, 2012; accepted September 19, 2012. This work has been supported by the Austrian Science Fund (FWF), proj-ects TRP102-N20 and S10503-N20; the European Regional Development Fund (EFRE) in the framework of the EU program Regio 13; the Chris-tian Doppler Research Association; the Federal Ministry of Economy, Family, and Youth; and the federal state of Upper Austria.

J. Bauer-Marschallinger, T. Berer, B. Reitinger, and P. Burgholzer are with Christian Doppler Laboratory of Photoacoustic Imaging and Laser Ultrasonics, Linz, Austria, and with Research Center for Non-Destructive Testing GmbH (RECENDT), Linz, Austria (e-mail: j.bauer-marschallinger@recendt.at).

H. Grün and H. Roitner are with Research Center for Non-Destructive Testing GmbH (RECENDT), Linz, Austria.

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Section II provides an overview of the measurement system and a description of the piezoelectric transducers used and the application of laser ultrasound. The section ends with a detailed explanation of the setup for tissue and liquid measurements. The samples investigated and used for testing the system are described in Section III. Section IV speciﬁes the signal processing, the signal loss model, and the determination of the parameters describing the ultrasound attenuation of the samples. Section V presents results for subcutaneous fat of pig, human blood, and olive oil. The paper concludes by discussing the method and by comparing the results of the measurements with previ-ously published studies.

II. M S

A. Overview

Ultrasonic attenuation was determined by comparing transmission measurements (see Fig. 1) of investigated samples and distilled water. The unfocused ultrasound beam was generated either by piezoelectric transducers or by laser ultrasonic excitation. Both generation methods are described in detail and compared in the three follow-ing subsections. After travelfollow-ing through the sample/water path, the ultrasound waves were detected by a piezoelec-tric transducer. For each sample and conﬁguration of ul-trasound generation, a reference measurement was per-formed in water with the same emitter–detector geometry as that used for the sample measurements. All experi-ments were done at room temperature (22°C to 24°C) in an air-conditioned room.

B. Generation With Piezoelectric Transducer

Unfocused piezoelectric immersion transducers (V358-SU, Panametrics, Waltham, MA) with a beam diameter of 6.35 mm, center frequency of 50 MHz, and 40 MHz full-width at half-maximum (FWHM) were used to emit and detect ultrasound. To maximize the received signals, the transducers must be aligned very accurately. Fig. 2 shows the sensitivity of the signal strength to deviations of co-axial alignment for the range of emitter–detector distances used in this study. Deviation of parallel alignment leads also to a decreased bandwidth, as Fig. 3 shows. The be-havior shown also occurs for emitter–detector distances of 27 and 40.25 mm.

The transducers were controlled by a pulser/receiver (5073PR-40-E, Olympus NDT Inc., Waltham, MA; max. gain 39 dB, min. gain −49 dB). The output of the ampli-ﬁer is linear up to 1 V peak when terminated by 50 Ω. The receiver gain is set to yield a voltage peak well below that limit. Two settings of the energy parameter of the pulser were used for the measurements. They are denoted E1 and E2 throughout the paper. The energy of the sig-nals generated with the E2 setting is 2.25 times higher than for the E1 setting.

C. Generation With Laser Ultrasound

The second ultrasound excitation method we used is based on a laser ultrasonic technique [19], [20]. In this case, a short laser pulse illuminates an opaque target. De-pending on the pulse duration and energy, this leads to thermoelastic or ablative generation of ultrasound waves. In the ﬁrst case, the absorbed energy leads to an abrupt temperature rise at the surface of the target and, hence, to thermoelastic expansion of the illuminated volume and the emission of ultrasonic waves. If the absorbed light power is high enough, the surface is vaporized and the

Fig. 2. The root-mean-square ultrasound signal generated and received by piezoelectric transducers is very sensitive to (top) out-of-axis displace-ment and (bottom) deviations of parallel aligndisplace-ment of the transducers.

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-  .:  -     2633

resulting rebound of the expanding plasma causes the emission of ultrasound waves within the target. This is called ablative generation. The laser pulses were gener-ated by a frequency-doubled Nd:YAG laser at wavelength of 532 nm, with a pulse duration of 6 ns, a repetition rate of 20 Hz, and pulse energies between 12 and 65 mJ (5% standard deviation). The diameter of the laser beam was about 6 mm. Because ablative excitation may hamper the comparability of sample and references measurements be-cause of changes of the optical absorption rate and ge-ometry of the surface, the pulse energy was chosen to be below the limit for ablative laser-ultrasound generation. Two effects were checked to determine this energy limit of approximately 65 mJ. An immediate effect is a visible light flash emitted by the plasma. Moreover, the surface of the target is blackened if ablation occurs. The thermoelas-tically generated ultrasound signals show negligible varia-tion (<1%) within typical operavaria-tional time (<10 min) for measuring a sample.

As targets for the laser pulses, we examined steel, alu-minum, and silicon with varying thickness. A silicon wafer with [111]-crystallographic orientation and a thickness of 4 mm turned out to be most useful for our setup. Mea-surements of laser ultrasound pulses generated by a silicon target and detected by piezoelectric transducers showed higher amplitude and bandwidth compared with steel and aluminum targets. Moreover, ablative ultrasound genera-tion on silicon occurs at higher pulse energies compared with the other materials. The minimal pulse energy neces-sary for ablation depends on many factors, such as boiling point, mass density, heat conductivity coeﬃcient, light ab-sorption [22]. The light penetration depth, which is about one to two orders of magnitude larger for silicon compared with aluminum and steel, also plays an important role for this limit.

D. Comparison of Piezoelectric and Laser-Generated Ultrasound

The ultrasound signals generated by the Nd:YAG

la-ser have lower peak frequencies fp, higher bandwidth, and

more intensity in relation to pulses emitted by the

piezo-electric transducer. Table I lists fp of the reference signals

measured for varying distances of detector and emitter in water. The peak frequency of the piezoelectric gener-ated signals is given for both energy settings, E1 and E2. Because the laser pulse energy has very little inﬂuence on

fp, Table I lists the average value and standard deviation

σ. The distance for the excitation with the piezoelectric

transducer is 0.65 mm greater because of the sample hold-er (see Section II-E). The peak frequency decreases with increasing distance because the ultrasonic attenuation of water is proportional to the square of the frequency. The peak frequency for 6.15/6.8 mm is reduced by sub-optimal alignment of detector and emitter. Because this alignment error aﬀected reference and sample measure-ments, the measured data was used to determine acous-tic attenuation. Table II lists the FWHM and the signal strength relative to piezoelectric excitation E1. Because the FWHM values are virtually constant over the emitter– detector distance, the table lists the average values with standard deviation. The signal strength is normalized to E1 before averaging over the varying distances. Although the standard deviation is high, the values show clearly the much higher intensity of the laser-generated signals. The relative increase of the strength of the ultrasound signals is less than the corresponding relative increase of the laser pulse energy. This is caused by the limited bandwidth of the piezoelectric detector and the bandwidth of the laser-generated ultrasound signals, which becomes broader with increasing pulse energy (see Section VI-A).

E. Measurement of Tissue

To achieve reproducible and accurate ultrasonic trans-mission measurements of tissue samples with thicknesses in the range of about 1 to 40 mm, we designed a cage sys-tem according to the following requirements: as we have seen from preliminary experiments with unfocused piezo-electric transducers (see Section II-B), it is important to align emitter and receiver in parallel. Moreover, the rela-tive position of the ultrasound source and detector must remain constant while performing sample and correspond-ing reference measurements. A 3-D model of the result-ing cage-system is depicted in Fig. 4. It mainly consists of two aluminum flanges holding the piezoelectric trans-ducers. Two 120°-shifted, radially directed screws fix the transducers within the flange. Dowel bolts ensure parallel alignment of detector and emitter. Defined sample thick-ness and constant emitter–detector distance is achieved by distance bolts. Screws are used to connect the two halves. The tissue sample is clamped between two aperture disks (see also Fig. 5) with an opening diameter of 7 mm. The thickness of the sample should exceed the length of the

TABLE I. P F   R U S  D E–D D. Piezo Laser Distance (mm) fp,E1 (MHz) fp,E2 (MHz) Distance (mm) fp (MHz) σ (MHz) 6.8 43 39 6.15 27 — 8.3 47 47 7.65 30.5 0.6 11.3 46 41 10.65 27.5 1 21.3 36 — 20.65 25.3 1.7

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2634

distance bolts by only a small amount to hold the sample in place without applying considerable stress. The tissue samples tend to slightly expand into the opening of the aperture disk. By removing one piezoelectric transducer and inserting an opaque target, the same cache system can be used for laser ultrasonic excitation. A sectional view of the setup in the case of laser ultrasound emitter is shown in Fig. 5. One piezoelectric transducer is removed to make way for the pulse laser beam. A fraction of the silicon wafer is clamped between the flange and an aper-ture disk. The figure also shows the small volume between the sample and the piezoelectric transducer (white area in Fig. 5). Three channels (see Fig. 4) on the faces of each sample holder connect this volume with the surrounding and ensure that it fills with water.

F. Measurement of Liquid

The setup for the liquid measurements is shown in Fig. 6. Two piezoelectric transducers placed in an aluminum pipe create a sample chamber into which the sample liquid or distilled water is filled. All liquid measurements were performed with the E1 setting of the piezoelectric pulser/ receiver (see Section II-B). This configuration provides sufficient signal strength in respect to the attenuation of the investigated liquids and widest possible bandwidth. Pipes with different lengths provide five chamber lengths in the range of 3.7 to 83.7 mm. As in the tissue measure-ments, two 120°-shifted, radially directed screws provide a defined position of the transducers within the pipe. The

chamber must be drained carefully between reference and sample measurements.

III. E M

Table III lists all investigated samples with thickness xS

and standard deviation σxS.

A. Fat Tissue of Pig

We decided to investigate porcine fat tissue, because its acoustic properties are comparable to human fat and it is more easily accessible. All samples were taken from the same Ö-HYB pig (hybrid of {Edelschwein × Landrasse} × Pietrain) bought from a local slaughterhouse and mea-sured within a day after excision. The tissue was kept cool in a refrigerator (5°C to 8°C) except for the time of mea-surement. The approximately 20 × 20 mm large pieces of soft subcutaneous fat had nominal thicknesses of 1.5, 3, 6, and 20 mm and were cut out from one piece. The dermis and epidermis were removed immediately before the measurement. When the sample was clamped between the two aperture disks, an additional thickness of approxi-mately 0.2 mm resulting from swelling was observed. The small pressure applied to the sample causes the soft fat to spread into the opening of the aperture disk. The thick-ness of this swelling is estimated by a sliding caliper. The uncertainty of thickness is taken into account in the evalu-ation of the measurements (see Section IV-E-1). Sample F1P and F1L are actually the same piece of fat tissue, but

TABLE II. A B  S S   R U S. Excitation FWHM (MHz) σ (MHz) SignalRMS (normalized to E1) σ (normalized to E1) Piezo E1 34 0 1 — Piezo E2 36 0.1 1.5 0.14 Laser 21 mJ 35 1.5 1.9 0.74 Laser 37.5 mJ 40 0.9 3.5 1.1 Laser 51 mJ 44 0.8 4.3 1 Laser 66 mJ 46 1 5.4 1.4 FWHM = full-width at half-maximum.

All-Optical Photoacoustic Imaging with Fiber-Optic Mach-Zehnder Interferometers / submitted by Johannes Bauer-Marschallinger

All-Optical

Photoacoustic Imaging

with Fiber-Optic

Mach-Zehnder

Interferometers

Doctoral Thesis

Doktor der technischen Wissenschaften

EIDESSTATTLICHE ERKLÄRUNG

Table of Contents

Danksagung

Zusammenfassung

Abstract

1 Motivation

2 Introduction

2.1 Photoacoustic Imaging

2.2 Photoacoustic Imaging with Integrating Detectors

2.3 Fiber-Optic Realization of Integrating Line Detectors for PAI

3 Published Articles

3.1 TUFFC: Broadband High-Frequency Measurement of Ultrasonic

Attenuation of Tissues and Liquids

Broadband High-Frequency Measurement of

Ultrasonic Attenuation of Tissues and Liquids

M