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Real-time photoacoustic and ultrasonic dual-modality imaging system for early gastric cancer: Phantom and ex vivo studies
Optics Communications 426 (2018) 519–525
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Optics Communications journal homepage: www.elsevier.com/locate/optcom
Real-time photoacoustic and ultrasonic dual-modality imaging system for early gastric cancer: Phantom and ex vivo studies Yongping Lin a,b , Zhifang Li a , Zuoran Li a , Jianyong Cai a , Huaqin Wui a , Hui Li a, * a
Fujian Normal University, College of Photonic and Electronic Engineering, Fujian Provincial Key Laboratory of Photonic Technology, Key Laboratory of Optoelectronic Science and Technology for Medicine, Ministry of Education, Fuzhou, China b Xiamen University of Technology, School of Optoelectronic and Communication Engineering, Xiamen, China
ARTICLE
INFO
MSC: 00-01 99-00 Keywords: Early gastric cancer Photoacoustic imaging Dual-modality imaging 3D imaging
ABSTRACT Photoacoustic (PA) and ultrasonic (US) imaging are promising techniques for imaging biological tissue and diagnosing internal organs. This paper presents a real-time PA and US dual-modality imaging system for early gastric cancer (EGC) based on a clinical US transducer. It has the advantage that only the PA excitation is inside the body but the US imaging system is outside the body, which makes the system less invasive for EGC detection and will be a potential tool for clinical use. The first experiments produced real-time PA images of blood vessel phantoms. Cross-sectional 2D PA images and reconstructed 3D PA images of the blood vessel phantoms demonstrate the system can clearly identity tissue structure. The experiments show that the system has a high axial spatial resolution but a relatively low lateral spatial resolution. The second experiments combining US and PA imaging, an artificial slightly elevated tumor and three artificial superficial flat tumors embedded inside the pig stomach mucosa ex vivo, demonstrate that the system is able to detect EGC accurately.
1. Introduction Pulmonary, gastric, esophageal, and liver cancers are the most common incident cancers and were identified as leading causes of cancer death in China [1]. Moreover, gastric cancer is the fifth most common cancer and the third most common cause of cancer death worldwide [2]. Imaging technologies play an important role in the clinical management of cancer, including screening, diagnosis, therapy planning and monitoring. Imaging techniques and technologies that enable the detection of early gastric cancer (EGC) are therefore very important. EGCs can be missed or misdiagnosed even using high definition endoscopy [3]. Due to the low tracer accumulation in diffuse and mucinous tumor types, positron emission tomography and computed tomography have a low detection rates [4]. X. Yan et al. [5] completed the construction of an ideal MRI/optical dual-modality molecular probe to diagnose gastric cancer. However, the concentrations of Fe3 O4 required to diagnose gastric cancer are considered excessive. Y. B. Ji et al. [6] investigated the feasibility of THz time-domain reflectometry for the discrimination of EGC from the normal gastric region. But these studies required a reliable miniaturized THz endoscope, currently unavailable.
Ultrasonic (US) imaging is widely used due to its advantages such as portability and high imaging speed, especially for guided examination, surgery or other imaging. However, US specificity in cancer detection is limited by the overlapping acoustic characteristics of benign and malignant solid lesions [7]. Photoacoustic (PA) imaging is a promising technique that complements US and is able to distinguish benign from malignant tumors. It is also non-ionizing and offers highcontrast imaging of both the surgical tools and photoabsorbers that are often encountered in diagnostic and therapeutic techniques [8–10]. PA imaging systems are considerably less expensive than conventional MRI and CT imaging systems and show promising for real-time clinical applications. A small modification of current US scanners [11–13] results in a shared detector platform (an array detector) that facilitates a natural integration of PA and US imaging, creating a hybrid imaging technique that combines functional (PA) and structural (US) information, fit for clinical translation [14–16]. Recently, K. Daoudi et al. [17] presented a compact and ergonomically designed hand-held probe for PA/US imaging. C. Yoon et al. [18] demonstrated a dual mode PA/US imaging system using a pixel-based focusing method. N. Wu et al. [19] developed a prototype ocular imaging system that integrates opticalresolution PA microscopy and high-frequency US imaging. E. Filoux
*
Corresponding author. E-mail addresses: [email protected] (Y. Lin), [email protected] (Z. Li), [email protected] (Z. Li), [email protected] (J. Cai), [email protected] (H. Wui), [email protected] (H. Li). https://doi.org/10.1016/j.optcom.2018.05.087 Received 21 November 2017; Received in revised form 20 April 2018; Accepted 29 May 2018 0030-4018/© 2018 Published by Elsevier B.V.
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Fig. 2. Experimental system schematic. Fig. 1. Schematic of PA and US dual-modality imaging system for EGC.
532 nm, 6-ns pulse duration, and the 10 Hz repetition rate to pump a tunable optical parametric oscillator (Surelite OPO Plus, Continuum) at the wavelength of 811 nm. To ensure that the radiation levels are maintained below the maximum permissible exposure (MPE) imposed by the American National Standards Institute (ANSI) for human skin [25], the beam was sent to a frosted glass to enlarge the radiation area after exiting the OPO. The linear array US transducer (0.3 mm pitch, 128 elements, 5–14 MHz bandwidth, 2–9 cm depth range) was also placed in the water tank, facing the sample, so as to capture the outgoing acoustic signal. The sampling rate for data acquisition was 40 MHz. The transducer was held in position by a precise 𝑋𝑌 -stage motion controller (SC300-2B, Zolix) which displaced it for scanning with a step size of 0.1-mm along the axis 𝑥. The whole system is controlled by a clinical US computer (Sonix Tablet) with an attached data acquisition device (Sonix DAQ) both manufactured by Ultrasonix. The tablet controls the motion controller and the laser. The data from the transducer is acquired by the Sonix tablet. The Sonix tablet is able to reconstruct an US image from the RF data of US from the 128 elements, employing a conventional US dynamic receive-delay-and-sum algorithm. It also constructs an image from the PA data through an improved US dynamic receive-delay-andsum algorithm which is more efficient for PA data. The system is capable of scanning 150 images/s. This rate is restricted by the physical limitation of ultrasound speed in conjunction with the imaging sequence. The sample was placed 60 mm away the transducer, so in order to cover the range of the sample, a reception depth 80 mm is required. Given the sampling rate of 40 MHz, the speed of sound in soft tissue (1540 m/s), and the number of detection channels (128), the system required 6.65 ms to acquire a single complete image. However, the pulsed 𝑄-switched Nd:YAG laser has a 10 Hz repetition rate, limiting the effective image rate of the system to 10 images/s. This demonstrates the fact that it is possible to render cross-section 2D PA and US images in real-time. We are able to acquire data in various positions with the 𝑋𝑌 -stage motion controller, but this system takes around 1.5 s to displace the transducer by 1 mm to a new position and acquire the PA signals. 3D PA images can only be reconstructed at a slow rate, depending on how many layers are required.
et al. [20] presented a combined PA and US imaging system used to obtain high-quality, co-registered images of mouse-embryo anatomy and vasculature. Nevertheless, none of them was used in the field of EGC detection. LH Wang et al. [21] presented simultaneous photoacoustic and ultrasonic dual-mode endoscopy and showed its ability to image internal organs in vivo, thus illustrating its potential clinical application. However, they used a focused US transducer which included optical fiber. That means both their photoacoustic signal excites and captures are located inside the body. This paper describes a real-time PA and US dual-modality imaging system for EGC detection employing a clinical US system. A fiber will be inserted the stomach through the esophagus to excite the PA signal; meanwhile, a linear array US transducer captures the signal outside the body as shown in Fig. 1. Thus, only a thin fiber with some holding devices needs to reside inside the patient’s body. This reduces the patient’s discomfort and simplifies the surgery process. It is label-free and combines the good sensitivity and high contrast-to-noise ratio of PA imaging with the clear morphological features of US imaging in visualizing stomach lesions. It is able to obtain PA and US cross-sectional images in real-time. From the 2D PA and US images, it is possible to reconstruct 3D images by transducer scanning. To demonstrate the initial feasibility of the system, two blood vessel phantoms and an ex vivo pig stomach embedded with tumors were imaged. 2. Methods and materials When a short-pulsed laser beam irradiates the tissue, some of the light is absorbed by it. This causes a small transient temperature rise which is then further converted to a pressure rise via thermoelastic expansion. The pressure rise is propagated as an ultrasonic wave, which can be detected by an US transducer as shown in Fig. 1. The initial local pressure (𝑃0 ) generated by the PA effect can be described as 𝑃 0 = 𝛤 𝜇𝑎 𝐹 ,
(1)
where 𝛤 is the dimensionless Grüneisen coefficient; 𝜇𝑎 is the optical absorption coefficient and 𝐹 is the local laser fluence. Eq. (1) has been used to estimate the optical parameters of the tissue [22,23].
2.2. Phantom experiments To demonstrate the system performance, two blood vessel phantoms were fabricated from two transparent soft plastic tubes (I.D. 0.5 mm, O.D. 1 mm, Length 60 mm) filled with fresh pig blood. The phantoms were placed in a solution of 2.5 g agar powder in 100 ml distilled water as shown in Fig. 3. These two tubes were laid out in parallel curves in order to validate the reconstructed PA image. The transducer was aligned orthogonally to the blood vessel phantoms, in order to maximize the spatial resolution. We reconstructed a 3D image by collecting 400 2D PA images at steps of 0.1 mm. We then analyzed the 3D image to determine the axial and lateral resolutions of the system.
2.1. General setup The experimental system schematic is shown in Fig. 2. The sample was immobilized in the water tank, and irradiated by a laser. The PA signals were captured by an US transducer. The maximal penetration depth of human stomach mucosa is about 1.9 mm at a laser wavelength of 811 nm [24]. To obtain a wavelength of 811 nm, we used a pulsed 𝑄-switched Nd:YAG laser (Surelite I-10, Continuum) at a wavelength of 520
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Optics Communications 426 (2018) 519–525 mJ radiated energy density on the tissue was approximately 4 cm 2 ), both 2D cross-sectional PA and US images were taken. 600 slices with a step of 0.1 mm along the axis 𝑦 were scanned to form a 3D image.
3. Results 3.1. Resolution A delay curve of one data channel is shown in Fig. 5(a), which represents the pre-beam forming data. Employing an improved US dynamic receive-delay-and-sum algorithm, a cross-sectional 2D PA image of the blood vessel phantoms was reconstructed, as shown in Fig. 5(b). The top boundary, the blood and the bottom boundary of the blood vessel phantoms are clearly distinguishable, making the depth of the blood vessel phantoms easily detectable. The PA signal of the blood vessel phantoms on the right is a weaker than the one on the left, due to the slightly greater distance from the radiation center which diminishes the light exposure. By analyzing the image intensities, the −6 dB axial and lateral resolutions are shown in Fig. 5(c) and (d) respectively. The results show that the full widths at half maximum (FWHM) of the system are 0.51 mm and 2.23 mm, respectively. The axial spatial resolution is very precise since the tube inner diameter is 0.5 mm. However, the lateral spatial resolution is relatively low. The lateral resolution could be improved by a filtering algorithm but it is limited by the transducer’s performance. Fig. 5(e) shows a 2D PA image of the blood vessel phantoms. Two blood vessel phantoms can be observed very clearly. One of the blood vessel phantoms is a little dimmer than the other because it is situated somewhat deeper. A 3D PA image of the blood vessel phantoms is also shown in Fig. 5(f). Due to the relatively low lateral resolution, the blood vessel phantoms look wider than the real blood vessel phantoms look like. Fortunately, the structural information is acquired correctly.
Fig. 3. Photograph of two blood vessel phantoms.
The total radiated energy was 15 mJ at 811 nm over an area of 12 cm2 mJ (i.e. 1.25 cm 2 ). This was below the MPE imposed by ANSI for human skin. 2.3. Pig stomach ex vivo experiments The human stomach walls consist of a mucosa, submucosa, muscular layer and serosa. The mucosa and submucosa were considered as a single mucosa layer because they are similar tissues. The serosa is neglected because it is very thin and tumors do not lodge into it in early stages. According to the macroscopic classification of EGC [26,27], Type 0 tumors are subdivided by stage into Type 0-I (protruding), Type 0-II (superficial) and Type 0-III (excavated). In addition, Type 0-II consists of Type 0-IIa (superficial elevated), Type 0-IIb (superficial flat) and Type 0-IIc (superficial depressed). In this paper, we are focusing on Type 0-IIa and Type 0-IIb which are the most concerning stages of EGC. A slightly elevated tissue is not always a tumor and further investigation is needed. Also, a superficial elevated tumor is easy to miss. A superficial depressed tissue could be detected in the same way as the slightly elevated tissue. It is very hard for a physician to detect a superficial flat tumor. In contrast, protruding and excavated tumors are easier to notice due to their specific shape. In order to validate the performance of the proposed system for realtime PA and US dual-modality imaging, we conducted scans of an ex vivo pig stomach. The top-view of the pig stomach sample is shown in Fig. 4(a). The region of interest (ROI) is a rectangular area of the sample (blue box). In order to simulate gastric cancer of the Type 0-IIa a pig blood clot was used, which is a reasonable substitute for a tumor, for the purpose of the experiment because tumors have a lot of more blood vessels from normal tissue. First, the ROI was cut out of the pig stomach. Then a small rectangular shaped cavity was scooped inside out from the mucosa of the pig stomach, and filled with the pig blood clot. Finally, the pig blood clot was covered with a thin layer of pig stomach mucosa, sliced off from another region of the mucosa of the pig stomach. The thickness of the pig stomach mucosa layer was below 1 mm, as shown in Fig. 4(b), in order to ensure that the light is able to penetrate and excite the PA signals. Fig. 4(c) shows the resulting artificial slightly elevated tumor, embedded inside the pig stomach mucosa. By repeating the procedure, three further artificial Type 0-IIb superficial flat tumors, embedded inside the pig stomach mucosa, were fabricated as shown in Fig. 4(d). The tumors were spaced apart to ensure the precise separation of the PA signals. After securing the samples in the water tank, tumor 1 was about 7.5 mm wide and 8.0 mm long. Tumor 2 was 5.0 mm wide and 8.0 mm long and tumor 3 was 4.0 mm wide and 4.0 mm long. Their depths were all preset at around 4 mm. After calculation, the volume of tumors 1, 2 and 3 were approximately about 240 mm3 , 160 mm3 and 64 mm3 , respectively. The tumors were stretched because of the pressure on the soft tissue under the water. Using conventional irradiation, the
3.2. Pig stomach ex vivo imaging Fig. 6(a) shows a real-time 2D US image of the pig stomach. Various artifacts resulted from sound rebound (from the tank) and, in some cases, other interferences with the signal. The structure of the stomach wall is accurately depicted by the US imaging. The slightly elevated tissue is clearly observable, indicating a possible tumor. The real-time 2D PA images of the pig stomach were also obtained simultaneously. A slightly elevated tumor PA image, superimposed on the US image, is shown in Fig. 6(c). Through further examination of the PA image, the slightly elevated tissue could be confirmed as a tumor. Tumor 0 is rapidly detected by the US imaging. The slightly elevated tumor is easily observed in the real-time US image. We have enough reason to believe that the same would be true for slightly depressed tumors. On the other hand, the superficial flat tissue is not visible, as shown in Fig. 6(b), therefore no tumor could be detected by US means alone. Superficial flat tumor PA images of the pig stomach were obtained simultaneously. A superficial flat tumor PA image, superimposed on the US image, is shown in Fig. 6(d). By examining the combined PA and US image, the three hidden tumors are now detected. This illustrates the US imaging provides guidance, while the real-time PA imaging is employed to acquire details of the tumor, through its high-contrast imaging capabilities. Fig. 7(a)–(c) show the magnified 2D PA image of the three superficial tumors along axis 𝑦 through the centers of the tumors. The shapes of the tumors are irregular due to the water pressure on soft tissue. The maximum width of tumor 1 is 7.9 mm, the minimum is 7.1 mm. The maximum width of tumor 2 is 7.0 mm, the minimum is 3.9 mm. The maximum width of tumor 3 is 4.3 mm, the minimum is 2.3 mm. Fig. 7(d) shows the magnified 2D PA image of the tumors along lateral direction through the center of tumors. Tumors 1, 2 and 3 are presented separately, according to the tumors placement. The maximum width of tumor 1 is 12.6 mm, the minimum is 9.8 mm and the height is 521
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Fig. 4. (a) Top-view of pig stomach sample. (b) Thin pig stomach mucosa layer. (c) An artificial slightly elevated tumor and (d) three artificial superficial flat tumors embedded inside the pig stomach mucosa.
Fig. 5. (a) One data channel of two blood vessel phantoms. Reconstructed 2D PA (b) and cross-sectional 2D PA (c) image of the blood vessel phantoms, respectively. (d) System axial resolution. (e) System lateral resolution. (f) 3D PA image of the blood vessel phantoms.
2.8 mm. This is different from the original size because the tumor was
much and their minimum widths are 5.8 mm and 4.1 mm, respectively.
vertically compressed. Tumor 2 and Tumor 3 were not compressed as
After calculation, taking the average maximum and minimum widths, 522
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Fig. 6. A real-time 2D US image of the slightly elevated tissue (a) and flat tissue (b) of the pig stomach, respectively. Superimposed PA images of the slightly elevated tumor (c) and three superficial flat tumors (d), respectively.
Fig. 7. Magnified 2D PA images of Tumor 1 (a) Tumor 2 (b) and Tumor 3 (c) along axis 𝑦 through center of tumors along lateral direction. Magnified 2D PA images (d) of the tumors along lateral direction through center of tumors along axis 𝑦. Reconstructed 3D (e) PA images of tumors.
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the volumes of the tumors are about 235 mm3 , 166 mm3 and 66 mm3 , respectively, almost identical to the volumes of the tumors we produced. In summary, the artificial tumors were detected correctly. Some artifacts occurred when the light irradiated the fringe of the pig stomach. Fig. 7(e) shows a reconstructed 3D PA image of the three superficial flat tumors. The structural information of the tumors can also be observed properly.
References
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4. Discussions PA imaging is good for improving detection at early stages. PA and US dual-modality imaging are better than US imaging alone. The proposed system detects the tumor size quite accurately, even though they are undetectable with US imaging. The ex vivo pig stomach experiments show the most difficult situation of detecting EGC. The challenges are the tumor depth. Type 0-I tumors are in the early stage and therefore it is best to detect and operate on them. Fortunately Type 0-I tumors protrude above the mucosa surface more than 0.5 cm in height [27]. Even the Type 0-II tumors will be appropriate to detect. Otherwise, with the increasing of the depth of tumor invasion, the system will lose its advantages because the large tumor could be detected by other techniques, such as the high definition endoscopy. One of the limitations of the PA system is its lateral resolution. The reconstructed images of blood vessel phantoms in Fig. 5(f) are wider than the real vessel phantoms due to the relatively low lateral resolution of the system. The main reasons are the transducer pitch width and the transducer bandwidth. A better transducer will increase the lateral resolution but it would also increase the cost. We could obtain better results by improving the signal beamforming algorithm. Since the PA signals will propagate for few microseconds before we capture them, a compensation algorithm might be a feasible approach to improving the lateral resolution. Future work will investigate a compensation algorithm optimization. Some artifacts occurred when the light irradiated the border of the pig stomach in Fig. 7(a). In clinical use, this will not happen because the stomach has no border. To get a better experimental result, a bigger pig stomach block is recommended to eliminate the fringe artifacts. The depth of the tumor was not studied because the depth of the hand-made hole in the thin pig stomach is hard to control. Furthermore, the depths of the blood clots and the holes vary with the water pressure on soft tissue. A precise study of the tumor size will be conducted in the future when we find a solution to control soft tissue sizes under the water. The biggest challenge of implementing the system in clinical applications is the 𝑄-switched Nd:YAG laser which is bulky, expensive and requires an optical table. A nanosecond pulsed laser diode (PLD) would be a reasonable excitation source [28]. High frame rate of PA B-scan imaging is feasible with the PLD and clinical ultrasound imaging system. The PLD is less costly, portable, and easy to handle, having potential for clinical applications. Naturally, the limitations of the PLD implementation (such as lower signal-to-noise ratio due to its smaller laser pulse energy) will be investigated and addressed. 5. Conclusions In summary, we proposed a real-time PA and US dual-modality imaging system for EGC based on a linear array US transducer. This system is less invasive than other proposed PA/US systems because the US is external and only the PA is internal. It has potential for better clinical outcomes. The axial resolution of the system appears to be very good, but the lateral resolution is not yet as good as the PA/US systems. Improving the lateral resolution will be the primary focus of future work. Acknowledgments This work is supported by the National Science Foundation of China No. 61675043, and 81571726. The authors wish to thank Peter Andreae from Victoria University of Wellington for useful discussions. 524
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Real-time photoacoustic and ultrasonic dual-modality imaging system for early gastric cancer: Phantom and ex vivo studies Yongping Lin a,b , Zhifang Li a , Zuoran Li a , Jianyong Cai a , Huaqin Wui a , Hui Li a, * a
Fujian Normal University, College of Photonic and Electronic Engineering, Fujian Provincial Key Laboratory of Photonic Technology, Key Laboratory of Optoelectronic Science and Technology for Medicine, Ministry of Education, Fuzhou, China b Xiamen University of Technology, School of Optoelectronic and Communication Engineering, Xiamen, China
ARTICLE
INFO
MSC: 00-01 99-00 Keywords: Early gastric cancer Photoacoustic imaging Dual-modality imaging 3D imaging
ABSTRACT Photoacoustic (PA) and ultrasonic (US) imaging are promising techniques for imaging biological tissue and diagnosing internal organs. This paper presents a real-time PA and US dual-modality imaging system for early gastric cancer (EGC) based on a clinical US transducer. It has the advantage that only the PA excitation is inside the body but the US imaging system is outside the body, which makes the system less invasive for EGC detection and will be a potential tool for clinical use. The first experiments produced real-time PA images of blood vessel phantoms. Cross-sectional 2D PA images and reconstructed 3D PA images of the blood vessel phantoms demonstrate the system can clearly identity tissue structure. The experiments show that the system has a high axial spatial resolution but a relatively low lateral spatial resolution. The second experiments combining US and PA imaging, an artificial slightly elevated tumor and three artificial superficial flat tumors embedded inside the pig stomach mucosa ex vivo, demonstrate that the system is able to detect EGC accurately.
1. Introduction Pulmonary, gastric, esophageal, and liver cancers are the most common incident cancers and were identified as leading causes of cancer death in China [1]. Moreover, gastric cancer is the fifth most common cancer and the third most common cause of cancer death worldwide [2]. Imaging technologies play an important role in the clinical management of cancer, including screening, diagnosis, therapy planning and monitoring. Imaging techniques and technologies that enable the detection of early gastric cancer (EGC) are therefore very important. EGCs can be missed or misdiagnosed even using high definition endoscopy [3]. Due to the low tracer accumulation in diffuse and mucinous tumor types, positron emission tomography and computed tomography have a low detection rates [4]. X. Yan et al. [5] completed the construction of an ideal MRI/optical dual-modality molecular probe to diagnose gastric cancer. However, the concentrations of Fe3 O4 required to diagnose gastric cancer are considered excessive. Y. B. Ji et al. [6] investigated the feasibility of THz time-domain reflectometry for the discrimination of EGC from the normal gastric region. But these studies required a reliable miniaturized THz endoscope, currently unavailable.
Ultrasonic (US) imaging is widely used due to its advantages such as portability and high imaging speed, especially for guided examination, surgery or other imaging. However, US specificity in cancer detection is limited by the overlapping acoustic characteristics of benign and malignant solid lesions [7]. Photoacoustic (PA) imaging is a promising technique that complements US and is able to distinguish benign from malignant tumors. It is also non-ionizing and offers highcontrast imaging of both the surgical tools and photoabsorbers that are often encountered in diagnostic and therapeutic techniques [8–10]. PA imaging systems are considerably less expensive than conventional MRI and CT imaging systems and show promising for real-time clinical applications. A small modification of current US scanners [11–13] results in a shared detector platform (an array detector) that facilitates a natural integration of PA and US imaging, creating a hybrid imaging technique that combines functional (PA) and structural (US) information, fit for clinical translation [14–16]. Recently, K. Daoudi et al. [17] presented a compact and ergonomically designed hand-held probe for PA/US imaging. C. Yoon et al. [18] demonstrated a dual mode PA/US imaging system using a pixel-based focusing method. N. Wu et al. [19] developed a prototype ocular imaging system that integrates opticalresolution PA microscopy and high-frequency US imaging. E. Filoux
*
Corresponding author. E-mail addresses: [email protected] (Y. Lin), [email protected] (Z. Li), [email protected] (Z. Li), [email protected] (J. Cai), [email protected] (H. Wui), [email protected] (H. Li). https://doi.org/10.1016/j.optcom.2018.05.087 Received 21 November 2017; Received in revised form 20 April 2018; Accepted 29 May 2018 0030-4018/© 2018 Published by Elsevier B.V.
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Fig. 2. Experimental system schematic. Fig. 1. Schematic of PA and US dual-modality imaging system for EGC.
532 nm, 6-ns pulse duration, and the 10 Hz repetition rate to pump a tunable optical parametric oscillator (Surelite OPO Plus, Continuum) at the wavelength of 811 nm. To ensure that the radiation levels are maintained below the maximum permissible exposure (MPE) imposed by the American National Standards Institute (ANSI) for human skin [25], the beam was sent to a frosted glass to enlarge the radiation area after exiting the OPO. The linear array US transducer (0.3 mm pitch, 128 elements, 5–14 MHz bandwidth, 2–9 cm depth range) was also placed in the water tank, facing the sample, so as to capture the outgoing acoustic signal. The sampling rate for data acquisition was 40 MHz. The transducer was held in position by a precise 𝑋𝑌 -stage motion controller (SC300-2B, Zolix) which displaced it for scanning with a step size of 0.1-mm along the axis 𝑥. The whole system is controlled by a clinical US computer (Sonix Tablet) with an attached data acquisition device (Sonix DAQ) both manufactured by Ultrasonix. The tablet controls the motion controller and the laser. The data from the transducer is acquired by the Sonix tablet. The Sonix tablet is able to reconstruct an US image from the RF data of US from the 128 elements, employing a conventional US dynamic receive-delay-and-sum algorithm. It also constructs an image from the PA data through an improved US dynamic receive-delay-andsum algorithm which is more efficient for PA data. The system is capable of scanning 150 images/s. This rate is restricted by the physical limitation of ultrasound speed in conjunction with the imaging sequence. The sample was placed 60 mm away the transducer, so in order to cover the range of the sample, a reception depth 80 mm is required. Given the sampling rate of 40 MHz, the speed of sound in soft tissue (1540 m/s), and the number of detection channels (128), the system required 6.65 ms to acquire a single complete image. However, the pulsed 𝑄-switched Nd:YAG laser has a 10 Hz repetition rate, limiting the effective image rate of the system to 10 images/s. This demonstrates the fact that it is possible to render cross-section 2D PA and US images in real-time. We are able to acquire data in various positions with the 𝑋𝑌 -stage motion controller, but this system takes around 1.5 s to displace the transducer by 1 mm to a new position and acquire the PA signals. 3D PA images can only be reconstructed at a slow rate, depending on how many layers are required.
et al. [20] presented a combined PA and US imaging system used to obtain high-quality, co-registered images of mouse-embryo anatomy and vasculature. Nevertheless, none of them was used in the field of EGC detection. LH Wang et al. [21] presented simultaneous photoacoustic and ultrasonic dual-mode endoscopy and showed its ability to image internal organs in vivo, thus illustrating its potential clinical application. However, they used a focused US transducer which included optical fiber. That means both their photoacoustic signal excites and captures are located inside the body. This paper describes a real-time PA and US dual-modality imaging system for EGC detection employing a clinical US system. A fiber will be inserted the stomach through the esophagus to excite the PA signal; meanwhile, a linear array US transducer captures the signal outside the body as shown in Fig. 1. Thus, only a thin fiber with some holding devices needs to reside inside the patient’s body. This reduces the patient’s discomfort and simplifies the surgery process. It is label-free and combines the good sensitivity and high contrast-to-noise ratio of PA imaging with the clear morphological features of US imaging in visualizing stomach lesions. It is able to obtain PA and US cross-sectional images in real-time. From the 2D PA and US images, it is possible to reconstruct 3D images by transducer scanning. To demonstrate the initial feasibility of the system, two blood vessel phantoms and an ex vivo pig stomach embedded with tumors were imaged. 2. Methods and materials When a short-pulsed laser beam irradiates the tissue, some of the light is absorbed by it. This causes a small transient temperature rise which is then further converted to a pressure rise via thermoelastic expansion. The pressure rise is propagated as an ultrasonic wave, which can be detected by an US transducer as shown in Fig. 1. The initial local pressure (𝑃0 ) generated by the PA effect can be described as 𝑃 0 = 𝛤 𝜇𝑎 𝐹 ,
(1)
where 𝛤 is the dimensionless Grüneisen coefficient; 𝜇𝑎 is the optical absorption coefficient and 𝐹 is the local laser fluence. Eq. (1) has been used to estimate the optical parameters of the tissue [22,23].
2.2. Phantom experiments To demonstrate the system performance, two blood vessel phantoms were fabricated from two transparent soft plastic tubes (I.D. 0.5 mm, O.D. 1 mm, Length 60 mm) filled with fresh pig blood. The phantoms were placed in a solution of 2.5 g agar powder in 100 ml distilled water as shown in Fig. 3. These two tubes were laid out in parallel curves in order to validate the reconstructed PA image. The transducer was aligned orthogonally to the blood vessel phantoms, in order to maximize the spatial resolution. We reconstructed a 3D image by collecting 400 2D PA images at steps of 0.1 mm. We then analyzed the 3D image to determine the axial and lateral resolutions of the system.
2.1. General setup The experimental system schematic is shown in Fig. 2. The sample was immobilized in the water tank, and irradiated by a laser. The PA signals were captured by an US transducer. The maximal penetration depth of human stomach mucosa is about 1.9 mm at a laser wavelength of 811 nm [24]. To obtain a wavelength of 811 nm, we used a pulsed 𝑄-switched Nd:YAG laser (Surelite I-10, Continuum) at a wavelength of 520
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Optics Communications 426 (2018) 519–525 mJ radiated energy density on the tissue was approximately 4 cm 2 ), both 2D cross-sectional PA and US images were taken. 600 slices with a step of 0.1 mm along the axis 𝑦 were scanned to form a 3D image.
3. Results 3.1. Resolution A delay curve of one data channel is shown in Fig. 5(a), which represents the pre-beam forming data. Employing an improved US dynamic receive-delay-and-sum algorithm, a cross-sectional 2D PA image of the blood vessel phantoms was reconstructed, as shown in Fig. 5(b). The top boundary, the blood and the bottom boundary of the blood vessel phantoms are clearly distinguishable, making the depth of the blood vessel phantoms easily detectable. The PA signal of the blood vessel phantoms on the right is a weaker than the one on the left, due to the slightly greater distance from the radiation center which diminishes the light exposure. By analyzing the image intensities, the −6 dB axial and lateral resolutions are shown in Fig. 5(c) and (d) respectively. The results show that the full widths at half maximum (FWHM) of the system are 0.51 mm and 2.23 mm, respectively. The axial spatial resolution is very precise since the tube inner diameter is 0.5 mm. However, the lateral spatial resolution is relatively low. The lateral resolution could be improved by a filtering algorithm but it is limited by the transducer’s performance. Fig. 5(e) shows a 2D PA image of the blood vessel phantoms. Two blood vessel phantoms can be observed very clearly. One of the blood vessel phantoms is a little dimmer than the other because it is situated somewhat deeper. A 3D PA image of the blood vessel phantoms is also shown in Fig. 5(f). Due to the relatively low lateral resolution, the blood vessel phantoms look wider than the real blood vessel phantoms look like. Fortunately, the structural information is acquired correctly.
Fig. 3. Photograph of two blood vessel phantoms.
The total radiated energy was 15 mJ at 811 nm over an area of 12 cm2 mJ (i.e. 1.25 cm 2 ). This was below the MPE imposed by ANSI for human skin. 2.3. Pig stomach ex vivo experiments The human stomach walls consist of a mucosa, submucosa, muscular layer and serosa. The mucosa and submucosa were considered as a single mucosa layer because they are similar tissues. The serosa is neglected because it is very thin and tumors do not lodge into it in early stages. According to the macroscopic classification of EGC [26,27], Type 0 tumors are subdivided by stage into Type 0-I (protruding), Type 0-II (superficial) and Type 0-III (excavated). In addition, Type 0-II consists of Type 0-IIa (superficial elevated), Type 0-IIb (superficial flat) and Type 0-IIc (superficial depressed). In this paper, we are focusing on Type 0-IIa and Type 0-IIb which are the most concerning stages of EGC. A slightly elevated tissue is not always a tumor and further investigation is needed. Also, a superficial elevated tumor is easy to miss. A superficial depressed tissue could be detected in the same way as the slightly elevated tissue. It is very hard for a physician to detect a superficial flat tumor. In contrast, protruding and excavated tumors are easier to notice due to their specific shape. In order to validate the performance of the proposed system for realtime PA and US dual-modality imaging, we conducted scans of an ex vivo pig stomach. The top-view of the pig stomach sample is shown in Fig. 4(a). The region of interest (ROI) is a rectangular area of the sample (blue box). In order to simulate gastric cancer of the Type 0-IIa a pig blood clot was used, which is a reasonable substitute for a tumor, for the purpose of the experiment because tumors have a lot of more blood vessels from normal tissue. First, the ROI was cut out of the pig stomach. Then a small rectangular shaped cavity was scooped inside out from the mucosa of the pig stomach, and filled with the pig blood clot. Finally, the pig blood clot was covered with a thin layer of pig stomach mucosa, sliced off from another region of the mucosa of the pig stomach. The thickness of the pig stomach mucosa layer was below 1 mm, as shown in Fig. 4(b), in order to ensure that the light is able to penetrate and excite the PA signals. Fig. 4(c) shows the resulting artificial slightly elevated tumor, embedded inside the pig stomach mucosa. By repeating the procedure, three further artificial Type 0-IIb superficial flat tumors, embedded inside the pig stomach mucosa, were fabricated as shown in Fig. 4(d). The tumors were spaced apart to ensure the precise separation of the PA signals. After securing the samples in the water tank, tumor 1 was about 7.5 mm wide and 8.0 mm long. Tumor 2 was 5.0 mm wide and 8.0 mm long and tumor 3 was 4.0 mm wide and 4.0 mm long. Their depths were all preset at around 4 mm. After calculation, the volume of tumors 1, 2 and 3 were approximately about 240 mm3 , 160 mm3 and 64 mm3 , respectively. The tumors were stretched because of the pressure on the soft tissue under the water. Using conventional irradiation, the
3.2. Pig stomach ex vivo imaging Fig. 6(a) shows a real-time 2D US image of the pig stomach. Various artifacts resulted from sound rebound (from the tank) and, in some cases, other interferences with the signal. The structure of the stomach wall is accurately depicted by the US imaging. The slightly elevated tissue is clearly observable, indicating a possible tumor. The real-time 2D PA images of the pig stomach were also obtained simultaneously. A slightly elevated tumor PA image, superimposed on the US image, is shown in Fig. 6(c). Through further examination of the PA image, the slightly elevated tissue could be confirmed as a tumor. Tumor 0 is rapidly detected by the US imaging. The slightly elevated tumor is easily observed in the real-time US image. We have enough reason to believe that the same would be true for slightly depressed tumors. On the other hand, the superficial flat tissue is not visible, as shown in Fig. 6(b), therefore no tumor could be detected by US means alone. Superficial flat tumor PA images of the pig stomach were obtained simultaneously. A superficial flat tumor PA image, superimposed on the US image, is shown in Fig. 6(d). By examining the combined PA and US image, the three hidden tumors are now detected. This illustrates the US imaging provides guidance, while the real-time PA imaging is employed to acquire details of the tumor, through its high-contrast imaging capabilities. Fig. 7(a)–(c) show the magnified 2D PA image of the three superficial tumors along axis 𝑦 through the centers of the tumors. The shapes of the tumors are irregular due to the water pressure on soft tissue. The maximum width of tumor 1 is 7.9 mm, the minimum is 7.1 mm. The maximum width of tumor 2 is 7.0 mm, the minimum is 3.9 mm. The maximum width of tumor 3 is 4.3 mm, the minimum is 2.3 mm. Fig. 7(d) shows the magnified 2D PA image of the tumors along lateral direction through the center of tumors. Tumors 1, 2 and 3 are presented separately, according to the tumors placement. The maximum width of tumor 1 is 12.6 mm, the minimum is 9.8 mm and the height is 521
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Fig. 4. (a) Top-view of pig stomach sample. (b) Thin pig stomach mucosa layer. (c) An artificial slightly elevated tumor and (d) three artificial superficial flat tumors embedded inside the pig stomach mucosa.
Fig. 5. (a) One data channel of two blood vessel phantoms. Reconstructed 2D PA (b) and cross-sectional 2D PA (c) image of the blood vessel phantoms, respectively. (d) System axial resolution. (e) System lateral resolution. (f) 3D PA image of the blood vessel phantoms.
2.8 mm. This is different from the original size because the tumor was
much and their minimum widths are 5.8 mm and 4.1 mm, respectively.
vertically compressed. Tumor 2 and Tumor 3 were not compressed as
After calculation, taking the average maximum and minimum widths, 522
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Fig. 6. A real-time 2D US image of the slightly elevated tissue (a) and flat tissue (b) of the pig stomach, respectively. Superimposed PA images of the slightly elevated tumor (c) and three superficial flat tumors (d), respectively.
Fig. 7. Magnified 2D PA images of Tumor 1 (a) Tumor 2 (b) and Tumor 3 (c) along axis 𝑦 through center of tumors along lateral direction. Magnified 2D PA images (d) of the tumors along lateral direction through center of tumors along axis 𝑦. Reconstructed 3D (e) PA images of tumors.
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the volumes of the tumors are about 235 mm3 , 166 mm3 and 66 mm3 , respectively, almost identical to the volumes of the tumors we produced. In summary, the artificial tumors were detected correctly. Some artifacts occurred when the light irradiated the fringe of the pig stomach. Fig. 7(e) shows a reconstructed 3D PA image of the three superficial flat tumors. The structural information of the tumors can also be observed properly.
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4. Discussions PA imaging is good for improving detection at early stages. PA and US dual-modality imaging are better than US imaging alone. The proposed system detects the tumor size quite accurately, even though they are undetectable with US imaging. The ex vivo pig stomach experiments show the most difficult situation of detecting EGC. The challenges are the tumor depth. Type 0-I tumors are in the early stage and therefore it is best to detect and operate on them. Fortunately Type 0-I tumors protrude above the mucosa surface more than 0.5 cm in height [27]. Even the Type 0-II tumors will be appropriate to detect. Otherwise, with the increasing of the depth of tumor invasion, the system will lose its advantages because the large tumor could be detected by other techniques, such as the high definition endoscopy. One of the limitations of the PA system is its lateral resolution. The reconstructed images of blood vessel phantoms in Fig. 5(f) are wider than the real vessel phantoms due to the relatively low lateral resolution of the system. The main reasons are the transducer pitch width and the transducer bandwidth. A better transducer will increase the lateral resolution but it would also increase the cost. We could obtain better results by improving the signal beamforming algorithm. Since the PA signals will propagate for few microseconds before we capture them, a compensation algorithm might be a feasible approach to improving the lateral resolution. Future work will investigate a compensation algorithm optimization. Some artifacts occurred when the light irradiated the border of the pig stomach in Fig. 7(a). In clinical use, this will not happen because the stomach has no border. To get a better experimental result, a bigger pig stomach block is recommended to eliminate the fringe artifacts. The depth of the tumor was not studied because the depth of the hand-made hole in the thin pig stomach is hard to control. Furthermore, the depths of the blood clots and the holes vary with the water pressure on soft tissue. A precise study of the tumor size will be conducted in the future when we find a solution to control soft tissue sizes under the water. The biggest challenge of implementing the system in clinical applications is the 𝑄-switched Nd:YAG laser which is bulky, expensive and requires an optical table. A nanosecond pulsed laser diode (PLD) would be a reasonable excitation source [28]. High frame rate of PA B-scan imaging is feasible with the PLD and clinical ultrasound imaging system. The PLD is less costly, portable, and easy to handle, having potential for clinical applications. Naturally, the limitations of the PLD implementation (such as lower signal-to-noise ratio due to its smaller laser pulse energy) will be investigated and addressed. 5. Conclusions In summary, we proposed a real-time PA and US dual-modality imaging system for EGC based on a linear array US transducer. This system is less invasive than other proposed PA/US systems because the US is external and only the PA is internal. It has potential for better clinical outcomes. The axial resolution of the system appears to be very good, but the lateral resolution is not yet as good as the PA/US systems. Improving the lateral resolution will be the primary focus of future work. Acknowledgments This work is supported by the National Science Foundation of China No. 61675043, and 81571726. The authors wish to thank Peter Andreae from Victoria University of Wellington for useful discussions. 524
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