See discussions, stats, and author profiles for this publication at: https://www.researchgate.net/publication/224696252 1E-5 Synergy and Applications of Combined Ultrasound, Elasticity, and Photoacoustic Imaging (Invited) Conference Paper in Proceedings of the IEEE Ultrasonics Symposium · November 2006 DOI: 10.1109/ULTSYM.2006.114 · Source: IEEE Xplore CITATIONS 52 READS 721 10 authors, including: Some of the authors of this publication are also working on these related projects: Low-cost enabling technology of Photodynamic therapy View project Ultrasound based Ablation Monitoring View project Salavat R Aglyamov University of Houston 153 PUBLICATIONS 2,988 CITATIONS SEE PROFILE Andrei Karpiouk Georgia Institute of Technology 73 PUBLICATIONS 1,957 CITATIONS SEE PROFILE Srivalleesha Mallidi Massachusetts General Hospital 94 PUBLICATIONS 2,431 CITATIONS SEE PROFILE Shriram Sethuraman Philips 35 PUBLICATIONS 848 CITATIONS SEE PROFILE All content following this page was uploaded by Shriram Sethuraman on 20 May 2014. The user has requested enhancement of the downloaded file.
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See discussions, stats, and author profiles for this publication at: https://www.researchgate.net/publication/224696252
1E-5 Synergy and Applications of Combined Ultrasound, Elasticity, and
Photoacoustic Imaging (Invited)
Conference Paper in Proceedings of the IEEE Ultrasonics Symposium · November 2006
Synergy and Applications of Combined Ultrasound, Elasticity, and Photoacoustic Imaging
S.Y. Emelianov,1 S.R. Aglyamov, 1 A.B. Karpiouk, 1 S. Mallidi, 1 S. Park, 1 S. Sethuraman, 1 J. Shah, 1 R.W. Smalling, 2 J.M. Rubin,3 W.G. Scott4
1Department of Biomedical Engineering, University of Texas at Austin, Austin, TX 78712 USA 2Division of Cardiology, University of Texas Health Science Center at Houston, Houston, TX 77030 USA
3Department of Radiology, University of Michigan Medical School, Ann Arbor, MI 48109 USA 4Winprobe Corporation, North Palm Beach, FL 33408 USA
Abstract—An advanced in-vivo imaging technology; namely, combined ultrasound, elasticity and photoacoustic imaging, capable of visualizing both structural and functional properties of living tissue, is presented. This hybrid imaging technology is based on the fusion of the complementary imaging modalities and takes full advantage of the many synergistic features of these systems. To highlight fundamental differences and similarities between the imaging systems and to appreciate advantages and limitations of each imaging system, the basic physics of each imaging system is described. The experimental aspects of combined imaging including hardware, signal and image processing algorithms, etc. are presented. Noise and primary artifacts associated with each imaging modality and combined imaging system are analyzed, and techniques to increase and optimize contrast-to-noise and signal-to-noise ratios in the images are discussed. Finally, biomedical and clinical applications of the combined ultrasound, elasticity and photoacoustic imaging ranging from macroscopic to microscopic imaging of pathology are demonstrated and discussed.
I. INTRODUCTION The noninvasive and quantitative visualization of
morphological and physiological properties of tissue is desired in the field of biomedical imaging. We present an advanced in-vivo imaging technology; namely, combined ultrasound, elasticity and photoacoustic imaging, capable of visualizing both structural and functional properties of living tissue. This hybrid imaging technology is based on the fusion of the complementary imaging modalities and takes full advantage of the many synergistic features of these systems.
The objective of the current study was to investigate whether synergy of ultrasound, photoacoustic and elasticity imaging is possible. In this paper, we describe the analytical, numerical and experimental studies to evaluate the performance of the combined ultrasound, photoacoustic, and elasticity imaging. The experimental aspects of combined imaging including hardware, signal and image processing
algorithms, etc. are described. Finally, we demonstrate and discuss biomedical and clinical applications of the combined ultrasound, elasticity and photoacoustic imaging in interventional cardiology, tissue engineering, and cancer detection, diagnosis and therapy monitoring.
II. BACKGROUND
A. Ultrasound Imaging Ultrasound imaging is a relatively inexpensive, real-time
imaging modality that is widely used in clinical practice. Ultrasound is excellent for non-invasive imaging and diagnosis of various tissue abnormalities, and it is extensively used for evaluating the prostate, breast, kidneys, liver, pancreas, thyroid, heart, and blood vessels. Ultrasound is also used to guide fine needle, tissue biopsy to test for tissue pathology and to identify cancerous tissue. In addition, ultrasound imaging is heavily used in obstetric medicine.
During the past decade, there have been unprecedented activities towards improving the quality of ultrasound imaging. New digital signal processing techniques have led to better images with higher signal-to-noise ratio and tissue contrast. Advances in transducer technology (increased number of elements, higher frequency and bandwidth, 1.5-D arrays, etc.) and introduction of digital scanners have further improved the accuracy, contrast and resolution of ultrasound images, and reduced noise and image slice thickness. Advanced ultrasound techniques have been developed including harmonic imaging and imaging with contrast agents. Along with improvements in image quality, clinical applications of sonography have also been extended to new fields. Overall, ultrasound is now widely accepted for many diagnostic and clinical applications.
There are, however, some limitations of ultrasound imaging. Perhaps the most important one is the ability of ultrasound imaging to identify all abnormalities – there may be low or no contrast between the abnormality and the surrounding tissue in ultrasound images. This is due to the fact that bulk modulus variations in soft tissue – the primary contrast mechanism in current ultrasound imaging – are only a few percent (Fig. 1). Sonography is also an operator-dependent, subjective test – the more experience the operator has, the better the examination and the results.
Partial support from National Institutes of Health under EB004963, CA110079, CA109440, and HL068658 grants, and American Heart Association under 0655033Y grant is gratefully acknowledged.
Figure 1: Bulk and shear modulus of various soft tissues.
Nevertheless, ultrasound imaging can be greatly improved based on new technologies and engineering advances including new ultrasound-based imaging techniques.
B. Elasticity Imaging Pathological processes and tissue abnormalities are usually
associated with changes in mechanical properties of soft tissue [1-3]. The success of palpation as a diagnostic tool is evidence of this. Manual palpation uses the "hardness" or "elasticity" of tissue sensed manually. Hardness is directly related to the value of the shear or Young's elastic modulus. Even today, palpation is widely used by primary care physicians and medical specialists, in conjunction with imaging methods and laboratory measurements, as a screening procedure for hard masses in the breast and prostate. The efficacy of manual palpation, however, is limited to relatively large (on the order of 10 mm) superficial abnormalities, and the information obtained is inherently subjective. Nevertheless, differences in Young's or shear modulus (i.e., the quantitative measure related to the information obtained by palpation) between tissue types can be many orders of magnitude [3]. (For incompressible materials such as soft tissues, the Young’s modulus is simply three times greater than the shear modulus.)
Ultrasound imaging cannot directly provide information about tissue elasticity (i.e., Young’s or shear moduli). However, ultrasound imaging can be used to measure internal motion induced by external deformations, from which the elastic properties of the tissue can be reconstructed [4, 5]. Indeed, ultrasound-based elasticity imaging consists of evaluation of externally or internally induced internal tissue motion, measurement of strain tensor components and, finally, reconstruction of the spatial distribution of the elastic modulus using displacement and strain images. For elasticity imaging, ultrasound imaging is nearly ideal since signal phase can be exploited to sensitively track internal tissue motion.
The contrast in elasticity imaging is determined by variations of Young's (or shear) modulus. The Young's moduli in biological tissues span an enormous dynamic range – the differences between various soft tissues are about 5-6 orders of magnitude. Figure 1 shows rough estimates of elastic moduli of tissues. Thus potential contrast in elasticity imaging is enormous. Therefore, detection and differentiation of tissue pathologies, which differ in mechanical (but not necessarily acoustic) properties from the surrounding tissue, might be possible based on a measurement of Young's or shear modulus [1, 6].
C. Photoacoustic Imaging In addition to elasticity contrast between normal and
abnormal tissue, optical absorption of tissue could be another possible contrast mechanism to detect and differentiate various tissue pathologies.
Photoacoustic [7] (also known as optoacoustic [8] and, generally, thermoacoustic [9]) imaging relies on the absorption of electromagnetic energy, such as light, and the subsequent emission of an acoustic wave (lightning and thunder). The tissue is irradiated with nanosecond pulses of low energy laser light. Through the processes of optical absorption followed by thermoelastic expansion, broadband ultrasonic acoustic waves are generated within the irradiated volume. Finally, the acoustic waves, detected using ultrasound transducer, are used to form an image of the internal tissue structure. The received photoacoustic signal contains information about both position (time of flight) and strength of the optical absorber (amplitude of the signal) within the tissue.
The contrast in photoacoustic imaging is primarily determined by differences in optical absorption of various types of tissues. Indeed, the amplitude of the thermoelastic response of the tissue is proportional to the optical absorption, i.e., the stronger the absorption, the stronger the photoacoustic signal. The selection of an appropriate excitation wavelength is based on the absorption characteristics of the imaging target. The measurements of optical properties of tissues are quite variable and offer only an approximate guide to the optical behavior of tissues (Fig. 2).
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Figure 2: Optical absorption (µa) and extinction (µa+ µs) coefficients of various tissues [10].
However, several observations can be made from the typical absorption spectrum of several tissue types [10]. In the near-infrared regions, between 2000 and 3000 nm, water is the dominant absorber; the light penetration depth (the distance through tissue over which diffuse light decreases in fluence rate to 1/e or 37% of its initial value) varies from about 1 mm to 0.1 mm over this region. At the other end of the spectrum, in the ultraviolet region near 300 nm, the absorption depth is shallow, owing to absorption by cellular macromolecules. In the 400-600 nm range, absorption by blood (hemoglobin) is very strong. In the central region between 600-1300 nm, tissue absorption is modest while contrast between tissue components remains high. Within 700-1000 nm wavelength, the average optical penetration depth is on the order of tens of millimeters –
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therefore, this spectral range is very suitable for photoacoustic imaging of tissue.
The contrast mechanism in photoacoustic imaging offers the prospect of identifying both anatomical features and different functional activity of tissues that may be indistinguishable using other imaging modalities such as ultrasound, MRI or X-ray alone. For example, malignant tumors are associated with higher blood content and increased concentrations of hemoglobin [11, 12]. The increased concentration of strongly absorbing molecules (hemoglobin and other porphyrins) was shown to yield 2-6 fold optical contrast between tumors and normal tissues. Furthermore, malignant tumors have enhanced and noticeably hypoxic blood content. In contrast, benign tumors have a normal level of blood oxygenation. Since oxy- and deoxy-hemoglobin have distinct absorption spectra in the near infrared spectral range, the photoacoustic imaging using several optical wavelengths allows for the quantitative assessment of physiologic characteristics of tumors.
Generally, 5-10 ns long pulses are used in photoacoustic imaging – longer pulses do not satisfy the stress confinement criteria (e.g., continuous wave irradiation produces no photoacoustic signal), and shorter pulses produce weaker photoacoustic transients. The laser fluence is determined by the maximum permissible radiation for tissue. The 25-30 mJ/cm2 laser fluence of near-infrared irradiation is sufficient to deliver optical energy to deep tissues and to produce photoacoustic transients – this laser fluence represents safe level of laser irradiation of tissue in the wavelength range from 700-nm to 1100-nm defined by the American National Standards [13]. Therefore, photoacoustic level of pulsed laser energy does not produce any damage to the tissue, resulting in a negligible temperature increase at the surface and within the tissue [14].
D. Combined Imaging The rationale to combine ultrasound, photoacoustic and
elasticity imaging is based on multiple factors [15]. Fusion of these ultrasound-based techniques may result in an advanced imaging system capable of simultaneous morphological (ultrasound imaging) and physiological (photoacoustic and elasticity imaging) assessment of tissue – a desperately needed imaging tool needed in many clinical applications. The presence of a common detector (ultrasound transducer) leads to ease of integration – combined imaging approach would be relatively inexpensive. The prospect of improving the effectiveness of pathology detection and diagnosis makes the integration clinically significant.
III. COMBINED INTRAVASCULAR IMAGING The combined intravascular ultrasound (IVUS), elasticity
(IVE) and photoacoustic (IVPA) imaging represents a sophisticated in vivo imaging technology [15] capable of direct assessment of both functional and morphological properties of coronary atherosclerotic plaques [16, 17]. This technique may allow precise diagnosis, disease characterization and evaluation of the effects of treatment thus efficiently guiding interventions and reducing associated morbidity and mortality [18].
Indeed, IVUS imaging can provide insight into the extent and distribution of atherosclerotic plaque, allowing
characterization of vessel wall and plaque morphology but IVUS imaging alone cannot reliably differentiate soft or lipid-rich, i.e., vulnerable, plaques [19]. IVUS-based elasticity imaging aims to detect regions of increased strain (or decreased hardness) that are prone to rupture, thus improving the discrimination between lipid-rich and fibrous plaque, traditionally a limitation of standard IVUS. Photoacoustic imaging can detect composition and functional changes in tissues – the optical properties of the imaged tissue can be used to discriminate between different tissue/cell types. Furthermore, the photoacoustic signal can be resolved with spatial resolution on the order of a few tens of micrometers – a scale appropriate for imaging thin fibrous cap. However, IVPA imaging will greatly benefit if the functional properties of the plaques are visualized in context of the overall anatomical structures of the surrounding tissues, i.e., in context of IVUS images.
The combined IVUS/IVE/IVPA imaging system is presented in Fig. 3 [20, 21]. The core of the system is a 2.5F (0.83-mm diameter) catheter with a 40 MHz, single element, intravascular ultrasound transducer (Atlantis™ SR Plus IVUS imaging catheter, Boston Scientific, Inc.) Other major components of the system are ultrasound pulser/receiver (Panametrics, Inc.) interfaced with IVUS catheter, microprocessor unit (Pentium 4 PC) with A/D digitizer (Gage Applied Technologies, Inc.), an Nd:YAG Q-switched pulsed laser operating at 532 nm wavelength (New Wave Research, Inc.), and a custom-built, microprocessor-controlled scanning assembly to allow rotation of either catheter or sample. The Nd:YAG laser operating at a maximum pulse repetition frequency of 20 pulses per second was capable of providing a maximum energy of 24 mJ. The sample was irradiated from outside while the IVUS imaging catheter was positioned inside the lumen. To reduce laser fluence on the vessel to approximately 1 mJ/cm2, the laser beam was broadened using a ground glass optical diffuser.
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Figure 3: Experimental setup for IVUS, IVPA and IVE imaging.
All components of the system were interfaced and controlled using custom-designed Lab Windows (National Instruments, Inc.) application such that temporally consecutive, spatially concurrent IVUS, IVE and IVPA imaging modes can be executed. Specifically, the photoacoustic response of the tissue, followed by ultrasound pulse-echo signal, was recorded (Fig. 4) as the sample was rotated around the longitudinal axis. In IVE imaging, the phantom was pressurized while the IVUS frames were collected. Typically, one IVUS and IVPA image
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consisted of 250 beams, i.e., 250 records were acquired for one complete revolution of the sample. Once the data were acquired, the microprocessor unit was used for signal processing and image formation. Since both IVPA and IVUS signals were captured at each position, no additional spatial co-registration of images was required.
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To demonstrate the combined IVUS/IVE/IVPA imaging, experiments were performed using tissue-mimicking phantoms closely resembling acoustical, mechanical and optical properties of the tissue. The phantoms were constructed using 8% PVA solution with 0.4% of 15-µm diameter silica particles. The specific results presented in Fig. 5 were obtained from a modeled vessel, i.e., 100 mm long, 9.5 mm outer diameter, and 2.6 mm inner diameter cylindrical phantom. Inside of the wall, a 2.5-mm diameter inclusion with elevated optical absorption (0.07% of ultra-small graphite flakes) and elastic modulus (increased number of freeze/thaw cycles) was placed.
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Figure 5: Intravascular ultrasound (IVUS), strain (with palpogram), photoacoustic (IVPA) images of the phantom.
The IVUS image of the phantom does not display well the inclusion located at 4-5 o’clock. However, the inclusion can be easily identified in IVE image and IVPA image because of altered mechanical and optical properties. In IVE image, the inclusion appears as a dark (low strain) region both in the strain map and in the palpogram [22] signifying harder material. The palpogram is a colored ring on the outside of the image with the position of the inclusion highlighted. In IVPA image, the inclusion produces high-amplitude transient pressure waves indicating stronger optical absorption.
Since both IVUS, IVE and IVPA images were obtained simultaneously, all images can be combined and the inclusion can be visualized in the context of the underlying tissue. In combined image (Fig. 6), the ultrasound signal is displayed if photoacoustic signal is smaller than the user-defined threshold, and vise versa. In addition, palpogram of IVE image is also
displayed here. Note the excellent correlation between IVE and IVPA images outlining the location and extent of the inclusion in the structural context of the vessel wall.
Figure 6: Combined IVUS/IVE/IVPA image of the vessel phantom.
The results presented in Figs. 5-6 argue favorably that combined IVUS/IVE/IVPA imaging is possible using clinically available IVUS catheters [16]. Both elasticity (strain) and photoacoustic imaging could be transparently and inexpensively integrated with IVUS imaging system offering a conceptually and technically advanced imaging tool for the detection and characterization of vulnerable plaques.
To further illustrate the ability of the combined intravascular imaging system, the IVUS and IVPA imaging was perform on ex vivo arterial samples obtained from a normal rabbit and a 9 months, low cholesterol diet, atherosclerotic rabbit (Fig. 7) and compared with an immuno-histochemical analysis of the tissue samples [18].
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Figure 7: IVUS and IVPA images of the atherosclerotic vessel and associated histological sections are contrasted with the IVPA image of the normal vessel.
An atherosclerotic rabbit aorta was imaged using a 40 MHz IVUS imaging catheter and a pulsed laser operating at 532 nm. The 6.75-mm diameter IVUS image suggests the presence of a concentric plaque although the extent of the plaque is unclear. The IVPA image confirms the presence of the lesions where the plaques manifest themselves as darker regions with bright IVPA signals at the boundary. The strong IVPA signal is more likely related to the dense population of macrophages in this
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region. Indeed, the fairly uniform (radial and azimuthal) IVPA image (5.5 mm diameter) from the normal vessel does not suggest a presence of an atherosclerotic plaque within this vessel. The histopathological analysis verifies the imaging results. Plaques with resulting smaller lumen are shown in tissue section stained with hematoxylin and eosin (H&E) to characterize the general plaque morphology. The higher expression (dark brown) of RAM 11 antigen shows that most macrophages are superficial (close to the lumen). Section stained with picrosirius red shows collagen all along the plaques and suggests slightly more collagen in the deeper regions.
The images presented in Fig. 7 demonstrate that combined imaging may provide clinically relevant and useful information about the plaque.
IV. OPTIMAL OPTICAL WAVELENGTH The photoacoustic images presented in Figs. 5-7 were
obtained using 532 nm optical wavelength. Generally, biomedical photoacoustic imaging could be performed in a wide range of optical frequencies. However, specific constraints and limitations of a particular application determine the best optical wavelength for photoacoustic imaging. For example, clinical photoacoustic imaging is more likely to be performed in a 600 to 1100 nm wavelength range since the blood absorption is relatively small, there is a reasonable penetration of the light into tissue, and different tissue constituents have different magnitude of the optical absorption providing reasonable contrast between the tissues (Fig. 2). If photoacoustic imaging is used to assess tumor structure, tumor-induced angiogenesis, hemoglobin content and the ratio of oxygenated to deoxygenated blood, a 650 nm optical wavelength maybe utilized. Therefore, for each application there maybe an optimal optical wavelength where photoacoustic imaging at the desired depth has maximum contrast based on the differences in absorption coefficients.
However, it also may be possible to produce image of tissue with a contrast based on a wavelength dependence of the optical absorption coefficient. Indeed, photoacoustic imaging can provide even more information related to tissue composition and function if photoacoustic measurements are performed at several wavelengths since the wavelength-dependent behavior of various tissue components is an independent parameter and can be used as an image contrast mechanism.
This so called spectroscopic photoacoustic imaging approach, illustrated in Fig. 8, is based on post-processing of the IVPA images obtained at several different wavelengths of laser irradiation [18]. We have used a tunable laser source (Brilliant B OPO system, Opotek, Inc.) operating within 680-900 nm. Note that since the same cross-section of sample is imaged, the temporal/spatial characteristic of photoacoustic response will not change from one image to another as it is primarily determined by the size and position of the optical absorber contributing to the signal. What would change is the amplitude of the response since it depends on the value of optical absorption coefficient. Therefore, by analyzing the magnitude changes of the IVPA signal at every position of the image (three different pixels are pictorially shown in Fig. 8), an
image displaying the dependence of the optical absorption on the wavelength can be made.
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Figure 8: Spectroscopic photoacoustic imaging of tissue structure, composition and function.
It is anticipated that the spectroscopic photoacoustic image will help to identify the composition of the tissue. In the intravascular imaging example presented in Fig. 8, three regions with a relatively strong photoacoustic response were selected. The wavelength-dependent changes in the magnitude of the photoacoustic transients at these locations are shown in Fig. 9. This figure also shows the IVUS and 780-nm IVPA images of the atherosclerotic vessel. Clearly, not only the magnitude of the photoacoustic signal (i.e., optical absorption coefficient) is different at these locations but also the wavelength dependence of the optical absorption is dissimilar. Similar differences can be observed in optical absorption spectra presented in Fig. 2.
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Figure 9: Changes in magnitude of the IVPA signal with optical wavelength. The IVUS and IVPA images of the atherosclerotic vessel are also shown for reference.
Therefore, spectroscopic photoacoustic imaging can identify a single optical wavelength for optimal photoacoustic imaging (maximum contrast based on the differences in absorption coefficient), and to produce image of tissue structure/composition (contrast based on wavelength dependence of absorption coefficient). Note that resolution in individual photoacoustic images and in spectroscopic photoacoustic image does not depend on the wavelength of the laser irradiation.
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V. SPATIAL RESOLUTION Spatial resolution in combined imaging is determined by
the properties of the ultrasound and laser systems (primarily transducer characteristics and duration of the laser pulse), the characteristics of the photoacoustic transients, and signal and image processing algorithms.
In ultrasound imaging using the fixed focus, mechanically scanned transducers, the axial (daxial) and lateral (dlateral) resolutions can be analyzed using simplified expressions for axial and lateral resolutions at the focus of a spherical radiator:
BWcdaxial 2
1= , #fdlateral ⋅= λ ,
where λ is the average wavelength, c is the speed of sound and BW is the bandwidth of the transducer [23, 24]. Clearly, higher frequency bandwidth increases axial resolution, and lower f-number (f#) transducers have tighter beam-widths, thus increasing lateral resolution. The spatial resolution in strain and elasticity imaging is primarily limited by the resolution of the ultrasound imaging but it is then further degraded by the performance of the speckle tracking algorithm (e.g., kernel size and filtering to reduce peak hoping and other motion tracking artifacts) [25-27]. The resolution in photoacoustic imaging is dictated by the frequency characteristics of the photoacoustic transient wave (determined by the absorber size and the length of the laser pulse) and receiving ultrasound transducer [17].
The spatial resolutions of IVPA and IVUS imaging were compared by imaging a cross-section of a 10 µm diameter carbon fiber. This fiber absorbs light and reflects sound and, therefore, is an appropriate point source for photoacoustic and ultrasound imaging. The target was placed 3 mm away from the transducer and imaged by recording 125 A-lines of the photoacoustic signal followed by the ultrasound pulse-echo signal (see Fig. 4). These A-lines covered 90-degree sector. The 4.5 mm radius IVUS and IVPA sector images are shown in Fig. 10. Clearly, the axial (or radial) resolution is better in IVPA image while lateral (or azimuthal) resolution is superior in IVUS B-Scan. The magnified IVUS and IVPA inserts further highlight the differences between two imaging modes.
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Figure 10: IVUS and IVPA images of a 10 µm diameter carbon fiber.
The reason for the differences in spatial resolutions between ultrasound imaging and photoacoustic imaging is the difference in the spectral content of the IVUS and IVPA signals (Fig. 11). In photoacoustic imaging, if the laser pulse (τpulse) is short and it satisfies the stress confinement criterion (τpulse < h/c, where h is size of the optical absorber and c is the speed of sound), the resulting photoacoustic response is broadband in nature although it is affected by the finite duration of the laser pulse and finite bandwidth of the receiving ultrasound transducer. In pulse-echo ultrasound imaging, an acoustic pulse of finite bandwidth is first transmitted into tissue and the backscattered signal is then received, i.e., both transmit and receive characteristics of the ultrasound transducer narrow the frequency bandwidth of the ultrasound signal. Clearly, the frequency bandwidth of photoacoustic signal is higher and, therefore, the axial resolution in IVPA imaging is expected to be higher compared to the IVUS imaging. Indeed, from the spectra in Fig. 11 and images in Fig. 10, the axial resolution of in IVUS imaging is about 55 µm while the axial resolution in IVPA imaging is approximately 40 µm [20]. Generally, therefore, the axial resolution in photoacoustic imaging should be slightly better compared to ultrasound imaging.
Lateral resolution in ultrasound and photoacoustic imaging is diffraction limited and depends on the average wavelength of the acoustic wave. Since the bandwidth of the detected photoacoustic signal is centered at the lower frequency compared to ultrasound pulse-echo signal, the effective wavelength of photoacoustic transient is increased and the lateral resolution is decreased, correspondingly. In the images presented in Fig. 10, the azimuthal (or lateral) resolution in IVUS imaging is about 3.2 degrees. In contrast, the lateral resolution in IVPA imaging is 5.5 degrees. Again, this differences in the lateral resolution between IVUS and IVPA images acquired using single element, fixed focus ultrasound transducer are due to the differences in frequency content of photoacoustic and ultrasound signals detected by the same ultrasound transducer.
Overall, the improvement in the frequency bandwidth characteristics of the ultrasound transducer will result in enhanced spatial resolution in photoacoustic and ultrasound (and therefore, strain/elasticity) imaging. The laser pulse duration in photoacoustic imaging must satisfy the stress confinement criterion given the size of the smallest resolvable absorber.
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VI. COMBINED IMAGING USING TRANSDUCER ARRAY The main building block of the combined imaging system is
an ultrasound imaging system. Today, most of the modern ultrasound imaging systems utilize transducer arrays. In addition, clinical applications of the combined imaging require real-time capability of the imaging system [28, 29]. The block diagram of the combined, real-time, array-based ultrasound, photoacoustic and elasticity imaging system is presented in Fig. 12 [29]. Here, an array of transducers is used both to transmit the ultrasound pulses and to receive the backscattered ultrasound signals for grayscale and Doppler ultrasound imaging. In photoacoustic imaging, a short laser pulse is transmitted into the tissue using light guides, such as optical fibers, integrated with the ultrasound transducer. Finally, the most common and clinically appropriate implementation of elasticity imaging is that in which the transducer itself is used to apply surface force and deform tissue while the ultrasound frames are acquired to estimate the internal tissue motion.
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Figure 12: Block diagram of the combined, array-based ultrasound, photoacoustic and elasticity imaging system.
The same ultrasonic transducer is used for ultrasound and photoacoustic imaging. However, the beamforming approaches in photoacoustic and ultrasound imaging are different. Since the light is rapidly scattering in tissue, the laser pulse irradiates the entire volume of the imaged tissue at once, i.e., it is difficult to focus the laser beam inside the tissue. Therefore, unlike conventional ultrasound imaging that utilizes focused transmitted beam and dynamically focused received beam, the photoacoustic signals are detected by each element of the array simultaneously [30]. Therefore, the photoacoustic imaging is similar to explososcan [31] or ultrafast [25-27] ultrasonic imaging, i.e., the entire photoacoustic image is obtained within one laser pulse.
We have modeled core components of each imaging modality to understand the design constrains and requirements of the combined ultrasound, photoacoustic and elasticity imaging system [29]. A 10 mm diameter circular inclusion positioned in the middle of a 50 mm think homogeneous phantom was imaged using a 40 mm aperture, 128 element linear array of transducers operating at 5 MHz (Fig. 12). The hyperechoic lesion was twice harder and contained higher concentration of the optical absorbers compared to the background. The geometry of the phantom and imaging settings closely correspond to setup shown in Fig. 12 except that light was delivered from the bottom of the phantom.
Simulated ultrasound, photoacoustic and strain images of the phantom with a single circular inclusion are shown in Figure 13. These images, centered around the inclusion, cover 30 mm (lateral) by 36 mm (axial) region of the phantom.
Ultrasound Image Strain Image Photoacoustic Image
Figure 13: Ultrasound, strain and photoacoustic images of a modeled phantom with a single rounded inclusion having elevated ultrasound backscattering coefficient, shear modulus and optical absorption coefficient.
As expected, the circular inclusion appears brighter in the grayscale ultrasound image since the lesion had more scatterers compared to the background. Clearly, ultrasound imaging can depict the underlying anatomical structures – the circular shape of the inclusion can be easily identified in the image. In the strain image the harder inclusion exhibits less strain compared to the surrounding soft tissue – this closely approximates the axial strain distribution derived in the analytical model of the phantom deformation. However, strain induced decorrelation noise is apparent in this image. While inclusion is clearly seen in ultrasonic and strain images, the boundaries of the inclusion in photoacoustic image are obscure – the inclusion appears to be spread in a lateral direction such that it is almost impossible to recognize the circular shape of the lesion. In addition, the light was attenuated as it reached the top and, therefore, the amplitude of the photoacoustic signal gradually decreased from the bottom to the top. This effect can be clearly seen in the inclusion itself – the lower parts of the inclusion have stronger signal compared to the top of the lesion. Furthermore, since the inclusion has more absorbers, there is a shadow region above the inclusion. Nevertheless, the information obtained from each image is different, and the combination of these three imaging techniques complements each other and provides additional information needed for the reliable detection and diagnosis of tissue pathology.
(A) (B)
Figure 14: Photoacoustic images of a phantom with a circular, highly absorbing inclusion. The images were obtained with an array of a) one-wavelength wide elements; b) three-wavelength wide elements.
The artifacts presented in photoacoustic image in Fig. 13 are related to the nature of photoacoustic imaging – as laser beam irradiates the tissue, the light quickly spreads throughout the tissue due to the optical scattering, i.e., the laser beam instantly interrogates the entire volume of tissue and the
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photoacoustic transients are simultaneously produced everywhere. Therefore, the beamforming in photoacoustic imaging can only utilize receive focusing but it may cause the artifacts when the inclusion with elevated optical absorption is imaged. To reduce these artifacts, various aperture apodization functions (e.g. Hanning, Blackman, etc.) can be applied. However, the major improvement in photoacoustic imaging may come from the changes of the angular response of the array elements achieved through the manipulations of the effective transducer element size [29].
Photoacoustic images obtained with different widths (one wavelength versus three wavelengths) of acoustic elements in the array are presented in Fig. 14. Both images here cover 40 mm by 50 mm area. Artifacts around the inclusion are clearly seen in the first image. Indeed, using small size elements (each acoustic transducer in 128 element array was one wavelength wide), the photoacoustic signal from a strong optical absorber is detected by nearly all elements since small transducers are sensitive to ultrasound waves sent from a wide angular range of directions. To achieve narrower directivity or angular response, elements three times of the size (three-wavelength wide) were used, while the ultrasound beams were formed using a one-wavelength sampled array of these elements. Here the artifacts are noticeably reduced and the shadow produced by the inclusion is clearly visible. However, these improvements are at the expense of lateral resolution – the speckle spots are further elongated in a lateral direction compared to the previous image. Thus, there is a tradeoff between the reduction of the artifacts and the lateral resolution.
However, lateral resolution of photoacoustic imaging can be improved given the difference in spectral content of photoacoustic and ultrasound signals. Indeed, images in Fig. 14 were obtained using f-number equals to three – this f-number was selected based on the ultrasound imaging criteria (e.g., directivity of transducer elements given the frequency content of the ultrasound signal). Since the photoacoustic signals are broadband and are centered at the lower frequency compared to ultrasound signals (see Fig. 11), the lower f-number and, therefore, sharper focusing can be used in photoacoustic imaging. This will result in a tighter beam thus improving lateral resolution in photoacoustic imaging.
VII. COMBINED IMAGING AT MICROSCOPIC RESOLUTION The combined ultrasound, photoacoustic and elasticity
imaging can be extended to microscopic resolutions. The combined biomicroscopy may be a vital imaging tool in several clinical and biomedical applications including microsurgery, tissue engineering [32], transplantation, etc. Indeed, combined biomicroscopy should be capable of accurate visualization of both structural and functional changes in tissues, sequential monitoring of tissue adaptation and/or regeneration, and possible assistance of drug delivery and treatment planning.
The combined imaging at microscopic resolution was evaluated on tissue mimicking phantoms imaged with 25 MHz single element focused transducer [32]. The results were obtained using two similar PVA-based phantoms where each phantom had either harder inclusion or optically absorbing inclusion (Fig. 15). Both inclusions were 700 µm diameter cylinders. The elastic (Young’s modulus) properties of the
inclusion and phantom materials were manipulated by controlling the degree of PVA crosslinking using mechanical (freeze/thaw cycles) method. The optical absorption of the inclusion was elevated by adding 0.1% by weight of ultra-fine graphite particles.
Single element focused transducer
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Figure 15: Experimental setup of combined biomicroscope.
To image the 10 mm (height) by 50 mm (width) phantoms, a water tank with an optical window on the bottom (Fig. 15) was used. The laser beam irradiated the phantom from underneath while the transducer, roughly aligned with the laser beam, was located above the phantom. To obtain the B-scan and photoacoustic images of the phantom, the transducer was mechanically scanned over the desired region.
Figure 16: Ultrasound, photoacoustic and combined images of the phantom with a single circular inclusion.
The results of ultrasonic and photoacoustic imaging are presented in Fig. 16. The location of the inclusion is clearly depicted in the ultrasonic B-scan image (4 mm by 2.4 mm area), but the image cannot explicitly portray the structure of inclusion. In contrast, the photoacoustic image unmistakably identifies the margins of the inclusion. These results presented in Fig. 16 suggest that photoacoustic imaging of highly absorbing blood vessels may complement Doppler ultrasound or color flow imaging of small vessels.
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The ultrasound and elasticity images of the phantom are presented in Fig. 17. Both images cover 6.25 mm by 3.36 mm area. The ultrasound B-scan image visualizes the size and location of the inclusion. However, the information regarding the biomechanical properties (harder and softer regions) of the phantom cannot be inferred directly from the B-scan image, as ultrasound imaging is not sensitive to variations in Young’s modulus of tissue. Since the inclusion was harder than the background material, the inclusion was expected to have lower strain magnitude. Lower strain values in the harder inclusion compared to the softer background are evident in Fig. 17.
Furthermore, the photoacoustic images could be acquired during tissue deformation. In tissue engineering, for example, if neovasculature is imaged, this technique may be used to quantitatively measure blood re-perfusion, i.e., the change of the photoacoustic signal as the tissue is deformed and blood is slowly pushed out of the vessels. This is demonstrated using a tissue mimicking phantom with a 700 µm diameter model of a blood vessel filled with optically absorbing liquid [33]. As the phantom was deformed from the surface, the amplitude of the normalized photoacoustic signal from the vessel gradually and monotonically decreases (Fig. 18). Furthermore, the spectral signature of the photoacoustic response changes as well – the smaller the vessel, the higher the frequency of the photoacoustic transient. Therefore, photoacoustic imaging and sensing can be possibly used to quantitatively assess neovasculature of the tissue without any exogenous contrast agents.
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The high-resolution combined ultrasound, photoacoustic and elasticity imaging could be used in several in-vivo applications. For example, in skin transplantation, the skin patches are removed from one area of the body and transplanted to another area. New blood vessels begin growing from the recipient area into the transplanted skin within few days. During the adaptation period, the graft should be constantly monitored for good blood circulation. The combined imaging technique is an appropriate tool for non-invasive evaluation of the growth of transplanted tissues and detection of angiogenesis in the recipient areas. Other applications of the combined biomicroscopy may include tissue engineering [32], cancer research [34], diagnostic imaging and therapy monitoring, cellular imaging [34], small animal imaging, microsurgery, etc.
VIII. CONTRAST AGENTS Generally, contrast agents are playing an increasingly
important role in biomedical and clinical applications, and are further being investigated for innovative medical and industrial applications. In diagnostic ultrasound applications, for example, microbubbles are often used as ultrasound contrast agents to enhance image quality [35]. Recently, microbubbles are being evaluated for targeted imaging of inflammation, angiogenesis, and also as adjunct microdevices for molecular imaging [36, 37] and cancer treatment [38], enhanced targeted drug delivery and gene therapy. Methods for single molecule detection and molecular imaging based on microbubble formation and detection were recently proposed [39]. Bubbles are also associated with cavitation bioeffects in soft tissue [40, 41], tissue erosion [42], and shock wave lithotripsy [41].
In photoacoustic imaging, blood often serves as an endogenous optical contrast agent [33]. Among exogenous optical contrast agents, dyes (e.g., indocyanine green dye) and various forms of nanoparticles (nanospheres, nanoshells, nanorods, nanocresents, etc.) are available. Gold nanoparticles are especially attractive since they can be conjugated to antibodies for molecular and site specific targeting. Gold nanospheres absorbing light at 532 nm have been investigated for molecular specific photoacoustic imaging [34]. The absorption spectra of the gold nanoparticles can be shifted by varying their shapes and sizes. For example, nanoshells were recently introduced – these nanoparticles have strong and tunable (600-1000 nm) absorption peaks in their optical extinction spectra [43]. Gold nanoshells have been applied in both photothermal therapy [44] and photoacoustic tomography [45]. Nanorods with absorption peak in the 700-800 nm range have also been used in flow and site specific imaging [46].
IX. THERAPY PLANNING, GUIDING AND MONITORING Another important role for combined imaging may be in
planning and guiding the therapy and monitoring the treatment response of the pathology. For example, the success of cancer treatment depends on the selection of the appropriate therapy. Often, however, the effect of the specific treatment is not known a priori, and it is even more important to identify the response of the tumor to therapy as early as possible.
Thermal treatments (e.g., RF ablation, high intensity focused ultrasound, photothermal therapy) induce a temperature increase to treat human neoplasm. Such therapies are attractive to both patients and physicians and are suggested as alternatives to open surgery and other invasive techniques in the treatment of tumors. In photothermal therapy, tuned contrast agents are used to enhance optical absorption of tissue irradiated with a continuous wave laser and, therefore, to produce a selective temperature increase of up to 40ºC in a tumor volume. However, to ensure successful outcome of photothermal therapy, the tumor needs to be imaged before therapy [47]. During the therapeutic procedure, spatial and temporal behavior of temperature increase must be monitored to ensure tumor necrosis while protecting surrounding tissues. In addition, immediately and long after the therapy, the tumor needs to be evaluated for necrosis and resurgence.
Combined imaging can be used to plan, guide and monitor the outcome of the photothermal therapy [47]. In this approach,
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combined imaging is first used prior to surgery to identify size, location and functional activity (uptake of the optical contrast agent) of the tumor. Then the ultrasound frames are recorded continuously during therapy and temperature maps are generated by tracking temperature-induced motion of the speckle pattern [48]. Furthermore, elasticity imaging can indicate progression of photothermal treatment by quantifying the mechanical properties of tissue. Using ultrasound images captured during externally or internally applied deformations, strain images are recorded during and after treatment. Since healthy, cancerous and thermally coagulated cells have different mechanical properties, elasticity imaging can indicate photothermal treatment efficacy. Finally, combined imaging can monitor the therapy outcome over time.
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In addition to ultrasound measurements of the temperature, photoacoustic imaging can be used to monitor the temperature changes [49]. This is demonstrated in Fig. 19. In this experiment, the fresh tissue sample (i.e., rabbit spleen) was placed in the heated water bath and imaged both ultrasonically (25 MHz) and photoacoustically (532 nm) as the temperature was gradually increased from room temperature to about 40ºC. The penetration depth of 532 nm light in tissue is not significant and, therefore, the maximum of the photoacoustic response is near the irradiated boundary. Nevertheless, the photoacoustic response clearly increases with temperature – this is due to the temperature dependence of the Gruneisen parameter [49].
Results presented in Fig. 19 and our previous studies [47] indicate that ultrasound imaging, ultrasound-based thermal imaging, elasticity imaging and photoacoustic imaging can be used to plan, guide and monitor the outcome of photothermal therapy where tumor and surrounding tissue are irradiated with continuous wave laser. Furthermore, our recent studies [34] suggest that pulsed laser photothermal therapy can also be guided by combined imaging.
X. CONCLUSIONS Combined ultrasound, photoacoustic and elasticity imaging
is practical and feasible. First, based on both anatomical (morphology) and functional (activity) properties of the tissue, it is anticipated that combined ultrasound, photoacoustic and elasticity imaging may detect and differentiate different types
of pathologies prior to significant anatomical or biochemical changes. Second, high frequency, and hence high spatial resolution, combined imaging is possible in most cases. Third, ultrasound is routinely used in many clinical procedures and, therefore, all necessary pre-requisites for combined imaging are readily available. Patients may not be subjected to any additional procedures, and the examination time will not increase significantly to perform combined imaging. Fourth, the combined imaging technique is non-ionizing and there are no safety concerns. Finally, combined imaging can be relatively inexpensive and portable.
ACKNOWLEDGMENTS The authors would like to acknowledge helpful discussions
with Dr. Sokolov, Dr. Milner, Dr. Oh, and Mr. Larson of the University of Texas at Austin, Mr. Amirian of the University of Texas at Houston Medical Center, and Dr. Litovsky of the University of Alabama at Birmingham. The authors also would like to thank Boston Scientific, Inc. for their technical support.
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