Photoacoustic array imaging of calcifications: phantom study

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1 Photoacoustic array imaging of calcifications: phantom study Yao-You Cheng 1, Tsai-Chu Hsiao 2, Wan-Ting Tien 3, Shih-Bin Luo 3, De-Yi Chiou 3, Meng-Lin Li 1, * 1 Dept. of Electrical Engineering, National Tsing Hua University, Taiwan 2 Institute of Photonics Technologies, National Tsing Hua University, Taiwan 3 Electronics and Optoelectronics Research Laboratories, Industrial Technology Research Institute, Taiwan * mlli@ee.nthu.edu.tw ABSTRACT Breast calcification is one of the most important indicators for early breast cancer detection. In this study, based on a medical ultrasound array imaging platform, we attempt to develop a real-time and high penetration photoacoustic (PA) array imaging system for visualization of breast calcifications. Phantom studies were used to verify the imaging capability and penetration depth of the developed PA array system for calcification imaging. Intralipid gelatin phantoms with different-sized hydroxyapatite (HA) particles - major chemical composition of the breast calcification associated with malignant breast cancers - embedded were imaged. Laser at 750 nm was used for photoacoustic excitation and a custom-made 5-MHz photoacoustic array transducer with linear light guides was applied for photoacoustic signal detection. Experimental results demonstrated that this system is capable of calcification imaging of mm HA particles. For the 0.5-mm HA particles, the imaging contrast was about 34 db and the achievable penetration was 20 mm where the axial, lateral, and elevational resolution of this PA array imaging system is 0.39 mm, 0.38 mm, and 1.25 mm, respectively. The highest frame rate was 10 frames/sec limited by the laser pulse rate. Overall, our results demonstrate that it is promising for PA imaging as a real-time diagnosis and biopsy guidance tool of breast micro-calcifications outside mass lesion. Future work will focus on optimization of the photoacoustic transducer to further improve the penetration depth and development of photoacoustic and ultrasound dual-modal imaging to enhance the calcification imaging capability. Keywords: photoacoustic imaging, breast cancer, calcification 1. INTRODUCTION Calcification is one of the important indicators for early breast cancer detection [1]. Breast calcifications can be found in both benign and malignant lesions, and they can be the first and earliest sign of malignancy. X-ray mammography is the well-established gold standard method for identifying micro-calcifications. However, X-ray mammography is with ionizing radiation and thus there is inevitably carcinogenic risk. The use of ultrasound imaging is a powerful adjunct to X-ray mammography for imaging of breast lesions. Nonetheless, it is challenging for ultrasound imaging of microcalcifications. Speckle noise results in low ultrasound contrast between breast tissues and micro-calcification [2]. Sickles pointed out that calcification particles smaller than 2 mm are unreliably detected using ultrasound [3]. Although state-of-the-art ultrasound imaging technology with higher probe frequency and spatial compounding may improve the visibility of micro-calcifications, generally it is difficult to detect micro-calcifications smaller than 0.5 mm [4], typically when they are located inside echogenic, fibro-glandular breast tissues instead of anechoic mass lesion. Photoacoustic (PA) imaging is an emerging bio-photonic imaging technique that overcomes the resolution drawback of pure optical imaging due to overwhelming light scattering in biological tissues, and possesses the merit of both optics and ultrasound namely, high optical absorption contrast and sub-millimeter ultrasound resolution up to an imaging depth of few centimeters. PA signals are induced as a result of transient thermo-elastic expansion when Photons Plus Ultrasound: Imaging and Sensing 2012, edited by Alexander A. Oraevsky, Lihong V. Wang, Proc. of SPIE Vol. 8223, SPIE CCC code: /12/$18 doi: / Proc. of SPIE Vol

2 biological tissues absorb the pulsed laser energy. These signals are then detected by ultrasound transducers and reconstructed to form images representative of optical absorption distribution. PA imaging can potentially not only overcome the drawbacks of both mammography and ultrasound but also takes the advantages of wavelength selective high optical contrast and high ultrasonic resolution for detection of breast lesions and micro-calcificaitons. Preclinical studies on human breast cancer have shown that near-infrared (NIR) PA imaging can reveal vascular features suggestive of malignancy [5][6]. In our previous study, we also have demonstrated the feasibility of visualization of calcifications using PA imaing by a PA microscopy setup [7]. It is found that 0.5 mm micro-calcificaitons own comparable PA signal intensity with blood at NIR range which shows that PA imaging and biopsy guidance of breast micro-calcifications outside mass lesion is promising. In this study, we investigate the feasibility of real-time and deep-penetration PA micro-calcification imaging. Based on a medical ultrasound array imaging platform, we attempt to develop a real-time and high penetration PA array imaging system for visualization of breast calcifications. Phantom studies were used to verify the imaging capability and penetration depth of the developed PA array system for calcification imaging. 2. MATERIALS AND METHODS 2.1 Photoacoustic/Ultrasound Dual Modal Array Imaging System The setup of our developed photoacoustic/ultrasound dual modal array imaging system was shown in Fig. 1. An optical parametric oscillator (OPO) (Surlite OPO Plus, Continuum, USA) pumped by a frequency-tripled Q-switched Nd:YAG laser (Surlite II-10, Continuum, USA) was employed to provide laser pulses with a pulse width of ~4 ns and a pulse repetition rate of 10 Hz. Because of the relatively high penetration in tissue, laser light at NIR range was used. The NIR tunable range of the used OPO was about 750 nm to 850 nm where the output energy is stronger and relatively suitable for deep penetration imaging. According to our previous study [7], 750-nm laser light could induce stronger PA signals of hydroxyapatite (HA) particles major chemical composition of the breast calcification associated with malignant breast cancers and owned less interference from blood; therefore, 750-nm laser was used as PA excitation source in this study. From the trend of the PA spectrum of HA calcifications obtained in our previous study, shorter wavelength, e.g., 700 nm, might provide even stronger PA signals of HA calcifications. For clinical use, it was preferred to integrate the ultrasound array probe and fiber bundles for laser delivery as a unit. Therefore, the used photoacsoutic transducer was composed of an 5-MHz ultrasound array probe (AT5L40B, Broadsound corporation) and two linear light guides, and was hold on a fixture connected to an X-Y-Z 3D motorized stage. The laser was coupled into the light guides and split to two sides of the ultrasound array probe. The cross of light beams from the two linear light guides was confocal with the elevational focal depth of the used ultrasound array probe. The tilting angles of the two linear light guides were adjustable in order to maximize the imaging depth. The maximum laser exposure on the phantom surface is about 5 mj/cm 2, which is less than the ANSI safety limite at 750 nm. The PA and ultrasound array signal detection was done with a PC-based ultrasound array data acquisition system Verasonics V-1 (Verasonics Inc.). Image reconstruction was done on the control PC. Here delay-and-sum beamforming algorithm was employed. To obtain better image contrast, coherence weighting was applied in addition to delay-and-sum beamforming [8]. Without signal averaging, real-time imaging could be done with our PA/ultrasound dual modal imaging system. The achievable frame for PA B-mode imaging and PA/US coherent B-mode imaging was 10 frames/sec which was limited by the laser repetition rate. In addition to B-mode imaging, C-scan could be done by moving the photoacoustic transducer along the elevation direction with scanning range of 35 mm limited by the stage itself. At the depth of 20 mm, the maximum achievable axial and lateral resolution was 0.39 mm and 0.38 mm, respectively. The estimated elevation resolution was about 1.25 mm. Proc. of SPIE Vol

3 Figure 1. Schematic diagram of the developed photoacoustic/ultrasound dual modal array imaging system 2.2 Phantom Preparation According to the human tissue optics data [9][10], tissue-mimicking calcification phantoms were made from gelatin (gelatin, G2500-1KG, SIGMA) and intralipid (Lipovenoes 20%). The absorption coefficient and reduced optical scattering coefficient of the custom-made phantoms were 0.13 cm -1 and 8.9 cm -1, respectively. All the blood-loaded tubes and HA calcification particles are buried at the depth of 20 mm. The placement was shown in Fig. 2(a). Three blood-loaded tubes with 0.25-mm inner diameter and HA particles with about 1.5-mm, 0.5-mm and 0.3-mm size were placed at the same horizontal plane. Fig. 2(b) was the photograph of the custom-made tissue-mimicking phantom. In order to verify the object positions imaged by PA imaging with ultrasound, there were no acoustic scatterers added in our phantom. That is, the resultant ultrasound B-mode and C-scan images could be free of speckle noise; thus forming a good reference for PA imaging. (a) (b) Figure 2. Illustration of the custom-made tissue-mimicking optical phantom (a) Placement of HA calcifications and blood-loaded tubes at the plane at the depth of 20 mm. red lines and circles represent blood-loaded tubes and HA particles, respectively. (b) Photograph of the custom-made tissue-mimicking phantom Proc. of SPIE Vol

4 3. EXPERIMENTAL RESULTS Figure 3 shows the PA and ultrasound B-mode images captured by our PA/ultrasound dual modal array imaging system. These images are displayed on a 50-dB dynamic range. The top panels are the PA B-mode images and the bottom panels are the ultrasound B-mode image. B-mode images of HA calcification particles with size of 1.5 mm, 0.5 mm, 0.3mm are shown in the order from left to right. There is no speckle noise in the acquired ultrasound B-mode images because no acoustic scatterers are added into the phantom. Therefore, the blood-loaded tubes and HA particles are clearly shown in the ultrasound B-mode images. That is, the object positions imaged by PA imaging can be verified by ultrasound B- mode images. PA imaging clearly visualizes all the HA calcification particles and blood-loaded tube at the depth of 20 mm. The estimated contrast between the imaging objects and background are 39.0 db for 1.5-mm HA, 34 db for 0.5-mm HA, 11.5 db for 0.3-mm HA, and 50.1 db for blood-loaded tubes. Note that in this study, one hundred times signal averaging is performed to improve the imaging signal-to-noise ratio. Figure 3. PA and ultrasound B-mode images of the custom-made tissue-mimicking phantom with HA particles and blood-loaded tube embedded. (a), (b), (c): PA B-mode images of 1.5-mm, 0.5-mm, and 0.3-mm HA particles, respectively. (d), (e), (f): ultrasound B-mode images of 1.5-mm, 0.5-mm, and 0.3-mm HA particles, respectively Figures 3(a) and (b) show the projected PA and ultrasound C-scan images, respectively, in the depth ranging from 15 mm to 25 mm. These projected C-scan images are displayed on a 50-dB dynamic range. Fig. 3(c) is the photograph showing the placement of the embeded objects in the tissue-mimicking phantom at the depth of 20 mm. In comparison of Fig. 3(a) with Fig. 3(c), all the objects are clearly shown in their respective positions in the PA projected C-scan image. Results from the ultrasound projected C-scan image also validate those obtained from PA imaging. The results demonstrate that HA particles can generate comparable PA signal intensity to that of blood.the anisotropic resolution between lateral and elevation direction causes the dimension distortion of the imaged object. The object dimension are larger in elevation direction compared to that in lateral direction. Proc. of SPIE Vol

5 Fig. 4 The projected PA (a) and ultrasound (b) C-scan images in the depth range from 15 mm to 25 mm. (c) Photograph showing the placement of the embeded objects in the tissue-mimicking phantom. Note that in this photograph, the tubes are not loaded with blood. 4. CONCLUSIONS Our experimental results demonstate that the developed PA/ultrasound dual modal array imaging system is capable of calcification imaging of mm HA calcification particles. For the 0.5-mm HA particles, the imaging contrast was about 34 db and the achievable penetration was 20 mm where the axial, lateral, and elevational resolution of this PA array imaging system is 0.39 mm, 0.38 mm, and 1.25 mm, respectively. The penetration depth is about adequate for imaging of breast calcifications, most of which appear at the depth from 15 mm to 35 mm [1]. The penetration depth is limited by the laser energy available in our current laser system and the coupling efficiency of the optical fiber bundles. In addition, from the trend of the PA spectrum of HA calcifications obtained in our previous study[7], shorter wavelength, i.e., 700 nm, might provide even stronger PA signals of HA calcifications. That is, deeper imaging depth for micro-calcifications may be achievable with shorter excitation wavelength, e.g., 700 nm, which is not available in our current laser system. Future work will focus on optimization of the photoacoustic transducer to further improve the penetration depth and development of photoacoustic and ultrasound dual-modal imaging to enhance the calcification imaging capability. ACKNOWLEDGEMENT Y.-Y. Cheng and T.-C. Hsiao contributed equally to this work. We acknowledge funding support from the Electronics and Optoelectronics Research Laboratories, ITRI, Taiwan (9351AA5110 and A35AA5110), the National Science Council, Taiwan (NSC E MY3 and NSC E MY2), and National Tsing Hua University, Taiwan (Boost Program 99N2905E1). We also appreciated Dr. Ren-Jei Chung from National Taipei University of Technology, Taiwan, for providing the HA particles. Proc. of SPIE Vol

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