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A Simple Technique for Visualizing Ultrasound Fields Without Schlieren Optics
Ultrasound in Med. & Biol., Vol. 41, No. 7, pp. 2071–2081, 2015 Copyright Ó 2015 World Federation for Ultrasound in Medicine & Biology Printed in the USA. All rights reserved 0301-5629/$ - see front matter
http://dx.doi.org/10.1016/j.ultrasmedbio.2015.03.004
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Original Contribution A SIMPLE TECHNIQUE FOR VISUALIZING ULTRASOUND FIELDS WITHOUT SCHLIEREN OPTICS NOBUKI KUDO Graduate School of Information Science and Technology, Hokkaido University, Sapporo, Japan (Received 20 September 2014; revised 5 March 2015; in final form 5 March 2015)
Abstract—A simple technique designed for visualization of ultrasound fields without Schlieren optics is introduced. An optical system of direct shadowgraphy with diverging light, which consists of a point light source and a shadow screen, constituted the basic system, but the screen was replaced by focusing optics: a camera that makes a virtual screen at its focus plane. The proposed technique visualizes displacement of light deflected by ultrasound, and the use of focusing optics enables flexible settings of the virtual screen position and optical magnification. Insufficient sensitivity of shadowgraphy was overcome by elimination of non-deflecting light using image subtraction of shadowgrams taken with and without ultrasound exposure. A 1-MHz focused transducer for ultrasound therapy and a 20-MHz miniature transducer for intravascular imaging were used for experiments, and alternate pressure change in short-pulsed ultrasound was visualized, indicating the usefulness of the proposed technique for evaluation of medical ultrasound fields. (E-mail: [email protected]) Ó 2015 World Federation for Ultrasound in Medicine & Biology. Key Words: Ultrasound field, Visualization, Optical method, Schlieren technique, Shadowgraphy, Focused shadowgraphy, Focused ultrasound, High-frequency ultrasound.
became an important application of the Schlieren technique (Azuma et al. 2002; Charlebois and Pelton, 1995; Hanafy 1979, 1980; Hanafy and Zanelli 1991; Neighbors et al. 1995; Neumann and Ermert 2006a, 2006b; Schneider and Shung 1996; Settles 2001; Weight and Hayman 1978). Exposure to continuous or long-burst ultrasound generates a cyclic change in the refractive index in a transparent medium, and the field works as an optical phase grating. Collimated light that propagates through the medium with a homogeneous refractive index converges on the focus of the second lens, and this non-deflected light is eliminated by the optical stop placed at the focus. At the same time, light that enters the phase grating produces diffraction, and the higher-order diffracted light that does not converge on the position of the optical stop produces a Schlieren image. Shadowgraphy (Biele 2003; Biss et al. 2009; Hargather and Settles 2007; Kudo et al. 2004, 2006, 2010; Omura et al. 2011; Settles 2001; Settles and Grumstrup 2005; Shimazaki et al. 2012) is another technique for visualizing inhomogeneities in the refractive index. The basic principle underlying this technique is also light deflection in gradient refractive index fields; however, the deflected light is simply
INTRODUCTION The Schlieren technique, which originated in studies started by Hooke more than 300 y ago (Davies 1981; Rienitz 1975, 1997), is used to visualize inhomogeneities in transparent media. This technique, which is based on deflection of light in a gradient field of the refractive index, has been playing an important role in the visualization of various characteristics of transparent objects such as residual stress, fluid flows and acoustic and shock wave fields. A typical Schlieren system consists of a pair of lenses with a point light source and a knife edge placed at the conjugate foci of the first and second lenses, respectively. Sensitive detection of inhomogeneity in the refractive index is realized by eliminating non-deflecting light from images. A variety of modified optical systems have been proposed and used for various applications (Settles 2001). After the discovery of light diffraction by ultrasound (Debye and Sears 1932; Lucas and Biquard 1932), visualization of ultrasound fields generated in water
Address correspondence to: Nobuki Kudo, Graduate School of Information Science and Technology, Hokkaido University, N14W9, Kitaku, Sapporo 060-0814, Japan. E-mail: [email protected] 2071
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Fig. 1. Three optical systems for visualization of ultrasound fields: (a) Dual-field lens Schlieren technique. (b) Direct shadowgraphy with a diverging light source. (c) Focused shadowgraphy with a diverging light source. (d) Convergence and divergence of light rays by an ultrasound field.
projected on the screen. Light deflection causes a positional shift in light rays incident on the screen and results in generation of an inhomogeneous intensity distribution of light on the screen, that is, a shadow. A novel technique based on simple optics (Kudo et al. 2004, 2006, 2010) was used in this study for visualization of an ultrasound field in water. The technique does not use a Schlieren optical system; it employs an optical system for focused shadowgraphy (Biss et al. 2009; Hargather and Settles 2007; Settles 2001). Two shadowgrams are taken with and without ultrasound exposure, and sensitive visualization of ultrasound fields is realized by subtraction of the shadowgrams to eliminate non-deflected light. The technique was applied to visualization of ultrasound fields of a focused transducer and a high-frequency miniature trans-
ducer, and the usefulness of the technique for visualization of short-pulsed ultrasound fields was investigated. METHODS Principle Figure 1 illustrates three optical systems for visualization of ultrasound fields: (i) a basic Schlieren optics system with two field lenses and a knife edge, (ii) a direct shadowgraph system with a diverging light source and a screen, (iii) the optical system used in this study. Optical system iii is a variation of an optical system called focused shadowgraphy, in which the screen in optical system ii is replaced by a camera. In direct shadowgraphy, we observe a shadow of an ultrasound field projected on screen M; however, with optical system iii, we observe
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Fig. 2. Principle underlying the new focused shadowgraphy technique. Two shadowgrams were taken (a) with and (b) without ultrasound exposure. (c) Software image subtraction of the two shadowgrams yields an ultrasound field image.
a shadow generated on virtual screen M0 , which is assumed to be placed just behind the acoustic field. Light transmitted through a medium with a gradient refractive index will be refracted in the direction that would increase the index. Because the refractive index of water increases with increase in pressure (Waxler and Weir 1963), the positive half-cycle of the ultrasound pressure field converges collimated light and the negative half cycle diverges, as illustrated in Figure 1d. The intensity distribution displayed on virtual screen M0 essentially contains shape distortion because the technique visualizes displacement of incident light. The new technique requires two steps to visualize ultrasound fields. In the first step (Fig. 2a, b), two shadowgrams are obtained with and without ultrasound exposure. The shadowgram obtained with ultrasound exposure consists of light deflected and not deflected by the ultrasound field, whereas the shadowgram obtained without ultrasound exposure contains only nondeflected light.
In the second step, these two shadowgrams are subtracted to obtain an image that contains only acoustically deflected light (Fig. 2c). Elimination of non-deflected light improves the low contrast of the shadowgrams. The Schlieren technique uses modulation of light phase and resulting light diffraction, whereas, shadowgraphy uses light refraction and the resulting positional shift in the incident light ray on a screen. Therefore, light not deflected by an ultrasound field can be eliminated using simple image subtraction. Experimental system Figure 3 illustrates the experimental system developed to confirm the practical usefulness of the proposed technique. A short-pulse laser diode (SLDHG-81/(SP), Hamamatsu Photonics, Shizuoka, Japan), 0.11 ns in pulse width and 850 mW in peak optical power, was used as a light source. A charge-coupled device (CCD) camera with 768 3 512 pixels and 15-bit A/D conversion (CV04, Mutoh, Tokyo, Japan) was used with a macro lens
Fig. 3. Experimental system. The optics includes only a light source and a camera. CCD 5 charge-coupled device; US 5 ultrasound.
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(Micro Nikkor 55 mm, Nikon, Tokyo, Japan). The light source was placed about 1.5 m before an ultrasound field, and the camera with a 0.25- or 0.35-m-working-distance (WD) lens was placed behind the field. The depth of field (DOF) measured at each WD was 2.9 or 4.5 mm, respectively. A stroboscopic technique was used to visualize an instantaneous field image of traveling pulsed ultrasound. Ultrasound pulses were irradiated at the repetition frequency of several kHz with the camera shutter open, and each pulse was followed by the light pulse with a time delay that determines the position of the traveling ultrasound pulse. The delay was controlled by a delay generator (DG535, Stanford Research Systems, Sunnyvale, CA, USA). Controls for the experiments, including delay setting, exposures to ultrasound and light pulses and image acquisition and subtraction, were programmed using Labview software (National Instruments, Austin, TX, USA). A Plexiglass water bath 115 (W) 3 70 (D) 3 150 (H) mm3 was used for experiments, and an acoustic absorber 20 3 20 3 10 mm3 was placed on the bottom. A sewing needle was fixed on the absorber and used as a target on which ultrasound and optical foci coincided. A very slow convection flow generated by an acoustic radiation force was observed as the movement of a tiny bubble or dust at the ultrasound focus. The bath was filled with degassed, filtered water to prevent disturbance of shadow images. All experiments were carried out using optical and ultrasonic elements mounted on a vibration isolator. Ultrasound transducers Two types of single-element transducers were used for visualization experiments. One is a focused transducer designed for ultrasound therapy, and the other is a miniature transducer used as an intravascular ultrasound (IVUS) probe. The focused transducer, a laboratoryassembled transducer with air backing, was 50 mm in diameter and 70 mm in focal length. The transducer was driven by a three-cycle pulse 1 MHz in center frequency. An arbitrary function generator (AFG310, Tektronix, Beaverton, OR, USA) and a radiofrequency amplifier (UOD-WB-1000, NEC Tokin, Miyagi, Japan) were used for driving the transducer, and the pressure waveforms at the ultrasound focus were measured using a polyvinylidene difluoride membrane hydrophone (MHA500 B, NTR Systems, Seattle, WA, USA). Figure 4 illustrates the pressure waveforms generated using the transducer driven by pulses with five increasing driving voltage settings. Strong non-linear waveforms were observed at higher driving voltages and the maximum peak positive and negative pressures were 7.9 MPa and 21.3 MPa, respectively. An acoustic field of the same focusing transducer was also visualized using a Schlieren system (Mizojiri Optical, Tokyo, Japan). For
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Fig. 4. Pressure waveforms generated by the focused transducer. The transducer was driven by 1-MHz three-cycle pulses with five increasing driving voltage settings, and the pressure waveforms were measured at the transducer focus using a membrane hydrophone.
this experiment, the transducer was driven by burstultrasound pulses 40 cycles in duration and 60.5 MPa in peak positive and negative pressures. A flat-faced miniature transducer 0.8 mm 3 0.5 mm in size and 20 MHz in center frequency was also used for visualization. The transducer was driven by a pulser/ receiver (5900 PR, Panametics, Waltham, MA, USA). The peak positive and negative pressures measured using the membrane hydrophone were 60.3 MPa at a distance of 3 mm from the transducer surface. RESULTS Figure 5 illustrates the acoustic fields of the focused transducer visualized using (a) the new technique and (b) the Schlieren technique. In Figure 5(a) is an instantaneous acoustic field at the ultrasound focus at a delay of 45 ms after insonation. The transducer was driven by a threecycle pulse with peak positive and negative pressures of 0.7 and 20.5 MPa, respectively, and shadowgrams were acquired with the camera with a WD of 0.35 m at the virtual screen position z of 30 mm (camera side). As mentioned earlier under Principle, the positive- and negative-half cycles of an ultrasound field converge and diverge a collimated light beam. Therefore, intensities of the subtracted image behind positive- and negativepressure fields have positive values (red) and negative values (blue), respectively. The main lobe of the focused transducer surrounded by side lobes was visualized in the image. Figure 5(b) is a Schlieren image acquired with the same transducer, but driven by the 40-cycle burst pulses. Different from our technique, the image was obtained
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Fig. 5. Focused ultrasound fields visualized using two techniques: (a) Ultrasound field of short-pulsed ultrasound visualized using the proposed technique. (b) Ultrasound field of long-burst ultrasound visualized using the Schlieren technique. The ultrasound waves are propagating from top to bottom.
with no synchronization. Note that the alternating pressure change is seen in Figure 5(a), whereas only the envelope of the pulse is visualized in the images in Figure 5(b). Full widths at half-maximum (FWHMs)
evaluated from Figure 5 were 2.4 mm (a) and (b) 4.0 mm at the ultrasound focus, respectively. The FWHM measured using the membrane hydrophone with a pressure-sensitive area 0.5 mm in diameter was 2.9
Fig. 6. Change in visualized images depending on the virtual screen position. The images were acquired at positions z of (a) 220 mm, (b) 210 mm, (c) 0 mm (focus of the ultrasound transducer), (e) 10 mm, (f) 20 mm and (g) 30 mm. Peak positive and negative pressures were 1.3 and 20.7 MPa, respectively. The images were clipped at the area of the main lobe. (d, h) Intensity profiles of images (a–c) and (e–g), respectively, evaluated on the vertical centerlines of the images.
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Fig. 7. Change in visualized images depending on the virtual screen position. Peak positive and negative pressures were changed to 7.9 and 21.3 MPa, respectively. Images were acquired at positions z of (a) 220 mm, (b) 210 mm, (c) 0 mm, (e) 10 mm, (f) 20 mm and (g) 30 mm. (d, h) Intensity profiles of the images.
mm, indicating the superiority of the new shadowgraph technique in determining beam width. Figures 6 and 7 illustrate changes in ultrasound field images around the ultrasound focus with respect to a position z of a virtual screen. The images were taken at six different positions from z 5 220 mm (light source side) to z 5 30 mm (camera side). Position z was changed by shifting the camera position with fixed WD to maintain the same optical magnification. The ultrasound focus (z 5 0) was determined by finding the peak of the pulse-echo signal from the top of the sewing needle. The pulser/receiver was used for that purpose. The optical focus was made to coincide with the ultrasound focus by adjusting the camera focus on the needle top. The pixel size of images calculated from a pixel count contained in one wavelength was 25 3 25 mm2. In Figure 6 are images acquired under exposure to pressures of 1.3 and 20.7 MPa. The images changed sensitively depending on the position of the virtual screen. Light deflection occurs at the ultrasound focus (z 5 0 mm); however, some field images were also visualized at z 5 220 and 210 mm (Fig. 6a, b). This phenom-
enon is known as a virtual shadowgram (Settles 2001). At positions z $ 0 mm, displacement of deflected light projected on a virtual screen is calculated by multiplying the deflection angle by the distance between the ultrasound focus and the virtual screen. Therefore, the ultrasound field was hardly visualized at z 5 0 mm (ultrasound focus, Fig. 6c), and the contrast of visualized images increased with increase in z (Figs. 6e–g). In Figure 6(d) and 6(h) are intensity profiles of the images in Figure 6(a–c) and 6(e–g), respectively, evaluated on the vertical centerlines of the images. The curves acquired at z 5 210 and 220 mm had three sharp positive peaks followed by slightly negative flat regions, which were different from the pressure waveform used for the experiments (Fig. 4, 1.3/20.7 MPa). Brightness profiles at z 5 10, 20 and 30 mm were more similar to their pressure waveforms; however, they were fundamentally different from the pressure waveforms because the brightness profiles were determined by projection of light deflected by pressure gradients. In Figure 7 are field images visualized under higher peak pressures (7.9/21.3 MPa). In the regions of
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Fig. 8. Ultrasound fields of five different pressure settings visualized using the proposed technique. The positions of the virtual screen were selected to minimize image distortion under high-pressure conditions and to maximize image sensitivity under low-pressure conditions.
z , 0 mm, the discrepancies between image and pressure waveform became more apparent. In the region of z . 0 mm, the asymmetry of the positive and negative regions that reflects the difference in positive and negative pressure waveforms is emphasized. In Figure 8 are field images acquired under exposure to five different ultrasound pressures. The positions of the virtual screen were selected to minimize image distortion and brightness saturation under high-pressure conditions and to maximize image sensitivity under low-pressure conditions. The series of images indicates the importance of virtual screen position for determination of a visualized pressure range. High-frequency ultrasound fields irradiated from the miniature rectangular transducer were visualized to evaluate spatial resolution. Images in Figure 9 were acquired by simply increasing optical magnification of the macro lens by setting WD to 0.25 m. The position of the virtual screen was not clearly determined in this visualization because of the large beam width of the flat-faced trans-
ducer. The pixel size of the image was 8 3 8 mm2. Propagation of the field was visualized by taking a series of images at different time delays. Complex acoustic fields near the plane transducer were visualized at delay times from 0.25 to 0.55 ms, and acoustic side lobes were visualized at delay times of 0.80 and 1.20 ms. Generation of an edge wave that propagates in the horizontal direction (direction of transducer rotation axis) was also visualized. Supplemental Video 1 (see online version at http://dx. doi.org/10.1016/j.ultrasmedbio.2015.03.004) illustrates the traveling short pulse. DISCUSSION Difference in configurations of the optical systems The simple optical system that contains only a light source and camera frees the proposed technique from sensitive adjustments. Furthermore, shadowgraphy, which uses light refraction, does not requires light phase integrity. Therefore, the proposed technique does not
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Fig. 9. Acoustic field propagation of a miniature transducer for intravascular ultrasound. The transducer was 0.8 3 0.4 mm2 in size and 20 MHz in center frequency.
require high-quality lenses and optical windows, which are essential for an optical system and a water bath used for Schlieren imaging. The angular spacing qd of a Fraunhofer diffraction pattern generated by an ultrasound phase grating is given by the equation sin qd 5
l L
(1)
where L is the ultrasound wavelength, and l is the light wavelength. qd has the value of 3.3 3 1024 rad under the conditions L 5 1.5 mm (1 MHz) and l 5 0.5 mm and has no pressure dependence. Small aberration Schlieren lenses with a long focal length are essential for sharp separation of higher-order light from the diffraction pattern with small angular spacing. A camera is also used for Schlieren imaging; however, in this situation, its lens also should have a long WD and, thus, a large DOF, suggesting that Schlieren images have little dependence on focusing position of the lens. On the other hand, the refraction angle qr of light deflected in a pressure gradient field dp=dx is given by the equations (Beach et al. 1973; Willard 1949)
tan qr 5 Cn0 $
dp=dx $w n0
ðn0 21Þ n20 11:4n0 10:4 Cn0 5 ðn20 10:8n0 11Þrc2
(2) (3)
where n0 is the refractive index without ultrasound exposure, Cn0 is the piezo-optic coefficient of water at n0 , w is light path length across the ultrasound field and r and c are the density and sound speed of water, respectively. Then, the maximum angle of qr with exposure to a sinusoidal wave of 1 MPa in peak pressure and 1 MHz in frequency is calculated to be 2.4 3 1023 rad by substituting n0 5 1.33, r 5 1 3 103 kg/m3, c 5 1,500 m/s and w 5 5 mm. Further increase is expected if an ultrasound wave has high non-linearity with a steep pressure increase at its wavefront. This means that refraction angles deflected by typical diagnostic or therapeutic ultrasound fields are at least 10 times larger than the diffraction angles, indicating that only a small part of deflected light is used for Schlieren images. A larger deflection angle allows use of a camera lens with a short WD and small DOF, yielding an effective focusing effect for which images
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exhibit sensitive change depending on the position of the virtual screen and improved sensitivity of shadowgraphy.
Difference in images acquired using the two techniques The Schlieren technique uses a diffraction pattern generated by interference of light at the second lens focus plane, and non-deflected light is subtracted by spatial filtering to remove the light that comes to the zerothorder position. Exposure to continuous or long-burst ultrasound is required to achieve clear separation between the zeroth- and higher-order light components. Basically, the Schlieren technique visualizes a time-averaged image with no synchronization to ultrasound phase. Therefore, the envelope of pressure variation that is not sensitive to ultrasound phase is visualized. As illustrated in Figures 5–9, the proposed technique has two advantages over the basic Schlieren techniques: (i) the ability to visualize alternating pressure change, and (ii) the ability to visualize pulsed ultrasound fields. The first point is not a fundamental difference because the stroboscopic Schlieren technique (Azuma et al. 2002; Charlebois and Pelton, 1995; Hanafy and Zanelli 1991; Neumann and Ermert 2006a, 2006b; Schneider and Shung 1996; Weight and Hayman 1978) enables visualization of alternating pressure changes by using short light pulses synchronized with an ultrasound phase. On the other hand, the stroboscopic Schlieren technique does not improve visibility of a short-pulse ultrasound field because the short burst length blurs its diffraction pattern and makes the spatial filtering difficult. There is no difficulty in visualization of short-pulsed ultrasound fields with the proposed technique. Various types of ultrasound fields, from 10.3/20.3 MPa to 17.9/21.3 MPa in peak positive and negative pressures and 1 and 20 MHz in center frequency, were visualized, as illustrated in Figures 8 and 9. Furthermore, light deflection angle is determined not by the pressure amplitude, but by the spatial gradient of the pressure distribution. This means that the deflection angles caused by diagnostic ultrasound of lower pressure and shorter wavelength could have a range of deflection angles similar to that caused by therapeutic ultrasound. These considerations indicate the possibility of using the technique to visualize both diagnostic and therapeutic ultrasound fields. For diagnostic applications, characterization of a pulsed ultrasound field is important because it directly determines performance of diagnostic ultrasound equipment. For therapeutic applications that use pulsed ultrasound, for example, extracorporeal shock wave therapy (Chung and Wiley 2002) and sonoporation (Kudo et al. 2009), visualization of an individual pulse shape is also important because the shape determines its therapeutic effects.
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Advantages of use of shadowgraphy Elimination of non-deflected light by image subtraction is another advantage of the proposed technique because elimination by spatial filtering requires careful adjustment of the high-quality optical system. Several studies have indicated the usefulness of image subtraction in the Schlieren technique for elimination of the timeindependent background (Azuma et al. 2002; Neumann and Ermert 2006a, 2006b). However, this subtraction has a different meaning from ours, because these techniques use Schlieren images from which nondeflected light is already removed. An inhomogeneous intensity distribution of illumination light causes artifacts in visualized images. Normalization of a subtracted image by one of the images used for the subtraction is effective (Omura et al. 2011; Shimazaki et al. 2012) if intensity fluctuation inside the light deflection area is negligible. However, we did not carry out normalization because our illuminating light beam had a significant inhomogeneity in intensity caused by interference of coherent light emitted from the laser diode. Advantages of use of focused shadowgraphy Use of a macro lens with a smaller DOF is also key. This enables imaging using large deflection light, which is hardly used in the Schlieren technique, and achieves acquisition of focused shadowgrams that vary with the position of the virtual screen. A flexible setting of the screen position is important because the presence of a physical screen disturbs the acoustic fields beside the screen. As illustrated in the results (Figs. 6 and 7), screen position affects both the sensitivity and distortion of visualized field images, and there is a trade-off between them (Fig. 8). This means that ultrasound fields of a very wide pressure range from IVUS to high-intensity therapeutic ultrasound can be visualized using one optical system simply by changing the virtual screen position. High-frequency ultrasound fields can also be visualized by simply increasing the magnification of the focusing optics (Fig. 9, Supplemental Video 1). The pixel size of the images taken using the macro lens placed at a WD of 0.25 m was 8 3 8 mm2. The simple optics enables combination with optical microscopy. For example, the pixel size of 0.93 3 0.93 mm2 was achieved using a 310 objective lens (Kudo et al. 2010). Limitations Image distortion caused by light deflection is the major limitation. This means that the proposed technique cannot replace hydrophone measurements. Shadowgraphy visualizes the second spatial derivative of an inhomogeneous refractive index distribution (Settles 2001),
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and the distribution can be reconstructed by double integration, in principle. Yet, it is difficult in practice because a shadowgram essentially contains image distortion caused by light deflection. Use of focused shadowgraphy reduces the distortion by placing the virtual screen closer to the ultrasound focus; however, reconstruction is still difficult because a shadow projected on a physical screen (Miyasaka et al. 2014) is different from an image captured using the camera with a finite DOF. Further studies are ongoing on this point. Under the condition of higher ultrasound intensity, non-linear steepening causes a rapid increase in light deflection angle. Light rays with larger deflection angles are eliminated by the pupil of the camera lens, and information on the steep pressure change is lost from the visualized images. Non-linear steepening can be observed in the waveform in Figure 4 (p1/2 5 17.8 MPa, 21.3 MPa), and the spatial peak temporal average intensity in the pressure cycle containing the peak positive and negative pressures was 174 W/cm2. Further increase in ultrasound intensity leads to cavitation phenomena, and light scattering caused by the cavitation bubbles disturbs field visualization. Pitts et al. (1994) proposed techniques that can determine an ultrasound field more precisely by detecting a phase shift of light integrated over its path transmitted through an ultrasound field. Two techniques, the phase recovery technique (Pitts et al. 1994) and phase contrast technique (Harigane et al. 2013; Pitts et al. 2001), were used for phase detection. The phase recovery technique estimates the phase shift by an iterative calculation using a magnitude image acquired at a known distance. The phase contrast technique derives the phase shift by using a magnitude image acquired using a p/2 delay column placed at the position of the zeroth-order light. These techniques can reconstruct 3-D fields using a CT algorithm in the pressure range of tens to hundreds of kilopascals in which light deflection is negligible. However, use of largely deflected light in shadowgraphy makes 3-D reconstruction difficult. CONCLUSIONS A simple technique suitable for visualization of medical ultrasound fields was proposed. The principle underlying the technique was described in comparison with two conventional techniques, the Schlieren technique and shadowgraphy, and the advantages and disadvantages were discussed. Two types of acoustic fields, a focused field and a high-frequency field, of short-pulsed ultrasound were visualized to confirm the practical usefulness of the proposed technique. The experimental results indicated that alternating pressure change in traveling short-pulsed ultrasound fields can be visualized
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clearly. Sensitivity and spatial resolution can be controlled flexibly simply by changing the focus position and magnification of the camera lens, indicating the usefulness of the technique for visualization of various medical ultrasound fields from small high-frequency IVUS ultrasound fields to high-intensity therapeutic ultrasound fields. The new shadowgraph technique cannot replace hydrophone measurements; however, this simple technique is useful not only for engineering laboratories, but also for the teaching environment to help laboratory staff who are not familiar with ultrasound understand its nature. Acknowledgments—The author expresses sincere gratitude to the late Professor Katsuyuki Yamamoto for his scrupulous discussion that initiated the research. Supportive comments from Professor Koichi Shimizu were invaluable. The author is also grateful to Professor Shin-ichiro Umemura for permission to use his Schlieren system. This study was supported by JSPS KAKENHI Grant 23300182.
SUPPLEMENTARY DATA Supplementary data related to this article can be found at http:// dx.doi.org/10.1016/j.ultrasmedbio.2015.03.004
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Original Contribution A SIMPLE TECHNIQUE FOR VISUALIZING ULTRASOUND FIELDS WITHOUT SCHLIEREN OPTICS NOBUKI KUDO Graduate School of Information Science and Technology, Hokkaido University, Sapporo, Japan (Received 20 September 2014; revised 5 March 2015; in final form 5 March 2015)
Abstract—A simple technique designed for visualization of ultrasound fields without Schlieren optics is introduced. An optical system of direct shadowgraphy with diverging light, which consists of a point light source and a shadow screen, constituted the basic system, but the screen was replaced by focusing optics: a camera that makes a virtual screen at its focus plane. The proposed technique visualizes displacement of light deflected by ultrasound, and the use of focusing optics enables flexible settings of the virtual screen position and optical magnification. Insufficient sensitivity of shadowgraphy was overcome by elimination of non-deflecting light using image subtraction of shadowgrams taken with and without ultrasound exposure. A 1-MHz focused transducer for ultrasound therapy and a 20-MHz miniature transducer for intravascular imaging were used for experiments, and alternate pressure change in short-pulsed ultrasound was visualized, indicating the usefulness of the proposed technique for evaluation of medical ultrasound fields. (E-mail: [email protected]) Ó 2015 World Federation for Ultrasound in Medicine & Biology. Key Words: Ultrasound field, Visualization, Optical method, Schlieren technique, Shadowgraphy, Focused shadowgraphy, Focused ultrasound, High-frequency ultrasound.
became an important application of the Schlieren technique (Azuma et al. 2002; Charlebois and Pelton, 1995; Hanafy 1979, 1980; Hanafy and Zanelli 1991; Neighbors et al. 1995; Neumann and Ermert 2006a, 2006b; Schneider and Shung 1996; Settles 2001; Weight and Hayman 1978). Exposure to continuous or long-burst ultrasound generates a cyclic change in the refractive index in a transparent medium, and the field works as an optical phase grating. Collimated light that propagates through the medium with a homogeneous refractive index converges on the focus of the second lens, and this non-deflected light is eliminated by the optical stop placed at the focus. At the same time, light that enters the phase grating produces diffraction, and the higher-order diffracted light that does not converge on the position of the optical stop produces a Schlieren image. Shadowgraphy (Biele 2003; Biss et al. 2009; Hargather and Settles 2007; Kudo et al. 2004, 2006, 2010; Omura et al. 2011; Settles 2001; Settles and Grumstrup 2005; Shimazaki et al. 2012) is another technique for visualizing inhomogeneities in the refractive index. The basic principle underlying this technique is also light deflection in gradient refractive index fields; however, the deflected light is simply
INTRODUCTION The Schlieren technique, which originated in studies started by Hooke more than 300 y ago (Davies 1981; Rienitz 1975, 1997), is used to visualize inhomogeneities in transparent media. This technique, which is based on deflection of light in a gradient field of the refractive index, has been playing an important role in the visualization of various characteristics of transparent objects such as residual stress, fluid flows and acoustic and shock wave fields. A typical Schlieren system consists of a pair of lenses with a point light source and a knife edge placed at the conjugate foci of the first and second lenses, respectively. Sensitive detection of inhomogeneity in the refractive index is realized by eliminating non-deflecting light from images. A variety of modified optical systems have been proposed and used for various applications (Settles 2001). After the discovery of light diffraction by ultrasound (Debye and Sears 1932; Lucas and Biquard 1932), visualization of ultrasound fields generated in water
Address correspondence to: Nobuki Kudo, Graduate School of Information Science and Technology, Hokkaido University, N14W9, Kitaku, Sapporo 060-0814, Japan. E-mail: [email protected] 2071
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Fig. 1. Three optical systems for visualization of ultrasound fields: (a) Dual-field lens Schlieren technique. (b) Direct shadowgraphy with a diverging light source. (c) Focused shadowgraphy with a diverging light source. (d) Convergence and divergence of light rays by an ultrasound field.
projected on the screen. Light deflection causes a positional shift in light rays incident on the screen and results in generation of an inhomogeneous intensity distribution of light on the screen, that is, a shadow. A novel technique based on simple optics (Kudo et al. 2004, 2006, 2010) was used in this study for visualization of an ultrasound field in water. The technique does not use a Schlieren optical system; it employs an optical system for focused shadowgraphy (Biss et al. 2009; Hargather and Settles 2007; Settles 2001). Two shadowgrams are taken with and without ultrasound exposure, and sensitive visualization of ultrasound fields is realized by subtraction of the shadowgrams to eliminate non-deflected light. The technique was applied to visualization of ultrasound fields of a focused transducer and a high-frequency miniature trans-
ducer, and the usefulness of the technique for visualization of short-pulsed ultrasound fields was investigated. METHODS Principle Figure 1 illustrates three optical systems for visualization of ultrasound fields: (i) a basic Schlieren optics system with two field lenses and a knife edge, (ii) a direct shadowgraph system with a diverging light source and a screen, (iii) the optical system used in this study. Optical system iii is a variation of an optical system called focused shadowgraphy, in which the screen in optical system ii is replaced by a camera. In direct shadowgraphy, we observe a shadow of an ultrasound field projected on screen M; however, with optical system iii, we observe
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Fig. 2. Principle underlying the new focused shadowgraphy technique. Two shadowgrams were taken (a) with and (b) without ultrasound exposure. (c) Software image subtraction of the two shadowgrams yields an ultrasound field image.
a shadow generated on virtual screen M0 , which is assumed to be placed just behind the acoustic field. Light transmitted through a medium with a gradient refractive index will be refracted in the direction that would increase the index. Because the refractive index of water increases with increase in pressure (Waxler and Weir 1963), the positive half-cycle of the ultrasound pressure field converges collimated light and the negative half cycle diverges, as illustrated in Figure 1d. The intensity distribution displayed on virtual screen M0 essentially contains shape distortion because the technique visualizes displacement of incident light. The new technique requires two steps to visualize ultrasound fields. In the first step (Fig. 2a, b), two shadowgrams are obtained with and without ultrasound exposure. The shadowgram obtained with ultrasound exposure consists of light deflected and not deflected by the ultrasound field, whereas the shadowgram obtained without ultrasound exposure contains only nondeflected light.
In the second step, these two shadowgrams are subtracted to obtain an image that contains only acoustically deflected light (Fig. 2c). Elimination of non-deflected light improves the low contrast of the shadowgrams. The Schlieren technique uses modulation of light phase and resulting light diffraction, whereas, shadowgraphy uses light refraction and the resulting positional shift in the incident light ray on a screen. Therefore, light not deflected by an ultrasound field can be eliminated using simple image subtraction. Experimental system Figure 3 illustrates the experimental system developed to confirm the practical usefulness of the proposed technique. A short-pulse laser diode (SLDHG-81/(SP), Hamamatsu Photonics, Shizuoka, Japan), 0.11 ns in pulse width and 850 mW in peak optical power, was used as a light source. A charge-coupled device (CCD) camera with 768 3 512 pixels and 15-bit A/D conversion (CV04, Mutoh, Tokyo, Japan) was used with a macro lens
Fig. 3. Experimental system. The optics includes only a light source and a camera. CCD 5 charge-coupled device; US 5 ultrasound.
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(Micro Nikkor 55 mm, Nikon, Tokyo, Japan). The light source was placed about 1.5 m before an ultrasound field, and the camera with a 0.25- or 0.35-m-working-distance (WD) lens was placed behind the field. The depth of field (DOF) measured at each WD was 2.9 or 4.5 mm, respectively. A stroboscopic technique was used to visualize an instantaneous field image of traveling pulsed ultrasound. Ultrasound pulses were irradiated at the repetition frequency of several kHz with the camera shutter open, and each pulse was followed by the light pulse with a time delay that determines the position of the traveling ultrasound pulse. The delay was controlled by a delay generator (DG535, Stanford Research Systems, Sunnyvale, CA, USA). Controls for the experiments, including delay setting, exposures to ultrasound and light pulses and image acquisition and subtraction, were programmed using Labview software (National Instruments, Austin, TX, USA). A Plexiglass water bath 115 (W) 3 70 (D) 3 150 (H) mm3 was used for experiments, and an acoustic absorber 20 3 20 3 10 mm3 was placed on the bottom. A sewing needle was fixed on the absorber and used as a target on which ultrasound and optical foci coincided. A very slow convection flow generated by an acoustic radiation force was observed as the movement of a tiny bubble or dust at the ultrasound focus. The bath was filled with degassed, filtered water to prevent disturbance of shadow images. All experiments were carried out using optical and ultrasonic elements mounted on a vibration isolator. Ultrasound transducers Two types of single-element transducers were used for visualization experiments. One is a focused transducer designed for ultrasound therapy, and the other is a miniature transducer used as an intravascular ultrasound (IVUS) probe. The focused transducer, a laboratoryassembled transducer with air backing, was 50 mm in diameter and 70 mm in focal length. The transducer was driven by a three-cycle pulse 1 MHz in center frequency. An arbitrary function generator (AFG310, Tektronix, Beaverton, OR, USA) and a radiofrequency amplifier (UOD-WB-1000, NEC Tokin, Miyagi, Japan) were used for driving the transducer, and the pressure waveforms at the ultrasound focus were measured using a polyvinylidene difluoride membrane hydrophone (MHA500 B, NTR Systems, Seattle, WA, USA). Figure 4 illustrates the pressure waveforms generated using the transducer driven by pulses with five increasing driving voltage settings. Strong non-linear waveforms were observed at higher driving voltages and the maximum peak positive and negative pressures were 7.9 MPa and 21.3 MPa, respectively. An acoustic field of the same focusing transducer was also visualized using a Schlieren system (Mizojiri Optical, Tokyo, Japan). For
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Fig. 4. Pressure waveforms generated by the focused transducer. The transducer was driven by 1-MHz three-cycle pulses with five increasing driving voltage settings, and the pressure waveforms were measured at the transducer focus using a membrane hydrophone.
this experiment, the transducer was driven by burstultrasound pulses 40 cycles in duration and 60.5 MPa in peak positive and negative pressures. A flat-faced miniature transducer 0.8 mm 3 0.5 mm in size and 20 MHz in center frequency was also used for visualization. The transducer was driven by a pulser/ receiver (5900 PR, Panametics, Waltham, MA, USA). The peak positive and negative pressures measured using the membrane hydrophone were 60.3 MPa at a distance of 3 mm from the transducer surface. RESULTS Figure 5 illustrates the acoustic fields of the focused transducer visualized using (a) the new technique and (b) the Schlieren technique. In Figure 5(a) is an instantaneous acoustic field at the ultrasound focus at a delay of 45 ms after insonation. The transducer was driven by a threecycle pulse with peak positive and negative pressures of 0.7 and 20.5 MPa, respectively, and shadowgrams were acquired with the camera with a WD of 0.35 m at the virtual screen position z of 30 mm (camera side). As mentioned earlier under Principle, the positive- and negative-half cycles of an ultrasound field converge and diverge a collimated light beam. Therefore, intensities of the subtracted image behind positive- and negativepressure fields have positive values (red) and negative values (blue), respectively. The main lobe of the focused transducer surrounded by side lobes was visualized in the image. Figure 5(b) is a Schlieren image acquired with the same transducer, but driven by the 40-cycle burst pulses. Different from our technique, the image was obtained
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Fig. 5. Focused ultrasound fields visualized using two techniques: (a) Ultrasound field of short-pulsed ultrasound visualized using the proposed technique. (b) Ultrasound field of long-burst ultrasound visualized using the Schlieren technique. The ultrasound waves are propagating from top to bottom.
with no synchronization. Note that the alternating pressure change is seen in Figure 5(a), whereas only the envelope of the pulse is visualized in the images in Figure 5(b). Full widths at half-maximum (FWHMs)
evaluated from Figure 5 were 2.4 mm (a) and (b) 4.0 mm at the ultrasound focus, respectively. The FWHM measured using the membrane hydrophone with a pressure-sensitive area 0.5 mm in diameter was 2.9
Fig. 6. Change in visualized images depending on the virtual screen position. The images were acquired at positions z of (a) 220 mm, (b) 210 mm, (c) 0 mm (focus of the ultrasound transducer), (e) 10 mm, (f) 20 mm and (g) 30 mm. Peak positive and negative pressures were 1.3 and 20.7 MPa, respectively. The images were clipped at the area of the main lobe. (d, h) Intensity profiles of images (a–c) and (e–g), respectively, evaluated on the vertical centerlines of the images.
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Fig. 7. Change in visualized images depending on the virtual screen position. Peak positive and negative pressures were changed to 7.9 and 21.3 MPa, respectively. Images were acquired at positions z of (a) 220 mm, (b) 210 mm, (c) 0 mm, (e) 10 mm, (f) 20 mm and (g) 30 mm. (d, h) Intensity profiles of the images.
mm, indicating the superiority of the new shadowgraph technique in determining beam width. Figures 6 and 7 illustrate changes in ultrasound field images around the ultrasound focus with respect to a position z of a virtual screen. The images were taken at six different positions from z 5 220 mm (light source side) to z 5 30 mm (camera side). Position z was changed by shifting the camera position with fixed WD to maintain the same optical magnification. The ultrasound focus (z 5 0) was determined by finding the peak of the pulse-echo signal from the top of the sewing needle. The pulser/receiver was used for that purpose. The optical focus was made to coincide with the ultrasound focus by adjusting the camera focus on the needle top. The pixel size of images calculated from a pixel count contained in one wavelength was 25 3 25 mm2. In Figure 6 are images acquired under exposure to pressures of 1.3 and 20.7 MPa. The images changed sensitively depending on the position of the virtual screen. Light deflection occurs at the ultrasound focus (z 5 0 mm); however, some field images were also visualized at z 5 220 and 210 mm (Fig. 6a, b). This phenom-
enon is known as a virtual shadowgram (Settles 2001). At positions z $ 0 mm, displacement of deflected light projected on a virtual screen is calculated by multiplying the deflection angle by the distance between the ultrasound focus and the virtual screen. Therefore, the ultrasound field was hardly visualized at z 5 0 mm (ultrasound focus, Fig. 6c), and the contrast of visualized images increased with increase in z (Figs. 6e–g). In Figure 6(d) and 6(h) are intensity profiles of the images in Figure 6(a–c) and 6(e–g), respectively, evaluated on the vertical centerlines of the images. The curves acquired at z 5 210 and 220 mm had three sharp positive peaks followed by slightly negative flat regions, which were different from the pressure waveform used for the experiments (Fig. 4, 1.3/20.7 MPa). Brightness profiles at z 5 10, 20 and 30 mm were more similar to their pressure waveforms; however, they were fundamentally different from the pressure waveforms because the brightness profiles were determined by projection of light deflected by pressure gradients. In Figure 7 are field images visualized under higher peak pressures (7.9/21.3 MPa). In the regions of
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Fig. 8. Ultrasound fields of five different pressure settings visualized using the proposed technique. The positions of the virtual screen were selected to minimize image distortion under high-pressure conditions and to maximize image sensitivity under low-pressure conditions.
z , 0 mm, the discrepancies between image and pressure waveform became more apparent. In the region of z . 0 mm, the asymmetry of the positive and negative regions that reflects the difference in positive and negative pressure waveforms is emphasized. In Figure 8 are field images acquired under exposure to five different ultrasound pressures. The positions of the virtual screen were selected to minimize image distortion and brightness saturation under high-pressure conditions and to maximize image sensitivity under low-pressure conditions. The series of images indicates the importance of virtual screen position for determination of a visualized pressure range. High-frequency ultrasound fields irradiated from the miniature rectangular transducer were visualized to evaluate spatial resolution. Images in Figure 9 were acquired by simply increasing optical magnification of the macro lens by setting WD to 0.25 m. The position of the virtual screen was not clearly determined in this visualization because of the large beam width of the flat-faced trans-
ducer. The pixel size of the image was 8 3 8 mm2. Propagation of the field was visualized by taking a series of images at different time delays. Complex acoustic fields near the plane transducer were visualized at delay times from 0.25 to 0.55 ms, and acoustic side lobes were visualized at delay times of 0.80 and 1.20 ms. Generation of an edge wave that propagates in the horizontal direction (direction of transducer rotation axis) was also visualized. Supplemental Video 1 (see online version at http://dx. doi.org/10.1016/j.ultrasmedbio.2015.03.004) illustrates the traveling short pulse. DISCUSSION Difference in configurations of the optical systems The simple optical system that contains only a light source and camera frees the proposed technique from sensitive adjustments. Furthermore, shadowgraphy, which uses light refraction, does not requires light phase integrity. Therefore, the proposed technique does not
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Fig. 9. Acoustic field propagation of a miniature transducer for intravascular ultrasound. The transducer was 0.8 3 0.4 mm2 in size and 20 MHz in center frequency.
require high-quality lenses and optical windows, which are essential for an optical system and a water bath used for Schlieren imaging. The angular spacing qd of a Fraunhofer diffraction pattern generated by an ultrasound phase grating is given by the equation sin qd 5
l L
(1)
where L is the ultrasound wavelength, and l is the light wavelength. qd has the value of 3.3 3 1024 rad under the conditions L 5 1.5 mm (1 MHz) and l 5 0.5 mm and has no pressure dependence. Small aberration Schlieren lenses with a long focal length are essential for sharp separation of higher-order light from the diffraction pattern with small angular spacing. A camera is also used for Schlieren imaging; however, in this situation, its lens also should have a long WD and, thus, a large DOF, suggesting that Schlieren images have little dependence on focusing position of the lens. On the other hand, the refraction angle qr of light deflected in a pressure gradient field dp=dx is given by the equations (Beach et al. 1973; Willard 1949)
tan qr 5 Cn0 $
dp=dx $w n0
ðn0 21Þ n20 11:4n0 10:4 Cn0 5 ðn20 10:8n0 11Þrc2
(2) (3)
where n0 is the refractive index without ultrasound exposure, Cn0 is the piezo-optic coefficient of water at n0 , w is light path length across the ultrasound field and r and c are the density and sound speed of water, respectively. Then, the maximum angle of qr with exposure to a sinusoidal wave of 1 MPa in peak pressure and 1 MHz in frequency is calculated to be 2.4 3 1023 rad by substituting n0 5 1.33, r 5 1 3 103 kg/m3, c 5 1,500 m/s and w 5 5 mm. Further increase is expected if an ultrasound wave has high non-linearity with a steep pressure increase at its wavefront. This means that refraction angles deflected by typical diagnostic or therapeutic ultrasound fields are at least 10 times larger than the diffraction angles, indicating that only a small part of deflected light is used for Schlieren images. A larger deflection angle allows use of a camera lens with a short WD and small DOF, yielding an effective focusing effect for which images
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exhibit sensitive change depending on the position of the virtual screen and improved sensitivity of shadowgraphy.
Difference in images acquired using the two techniques The Schlieren technique uses a diffraction pattern generated by interference of light at the second lens focus plane, and non-deflected light is subtracted by spatial filtering to remove the light that comes to the zerothorder position. Exposure to continuous or long-burst ultrasound is required to achieve clear separation between the zeroth- and higher-order light components. Basically, the Schlieren technique visualizes a time-averaged image with no synchronization to ultrasound phase. Therefore, the envelope of pressure variation that is not sensitive to ultrasound phase is visualized. As illustrated in Figures 5–9, the proposed technique has two advantages over the basic Schlieren techniques: (i) the ability to visualize alternating pressure change, and (ii) the ability to visualize pulsed ultrasound fields. The first point is not a fundamental difference because the stroboscopic Schlieren technique (Azuma et al. 2002; Charlebois and Pelton, 1995; Hanafy and Zanelli 1991; Neumann and Ermert 2006a, 2006b; Schneider and Shung 1996; Weight and Hayman 1978) enables visualization of alternating pressure changes by using short light pulses synchronized with an ultrasound phase. On the other hand, the stroboscopic Schlieren technique does not improve visibility of a short-pulse ultrasound field because the short burst length blurs its diffraction pattern and makes the spatial filtering difficult. There is no difficulty in visualization of short-pulsed ultrasound fields with the proposed technique. Various types of ultrasound fields, from 10.3/20.3 MPa to 17.9/21.3 MPa in peak positive and negative pressures and 1 and 20 MHz in center frequency, were visualized, as illustrated in Figures 8 and 9. Furthermore, light deflection angle is determined not by the pressure amplitude, but by the spatial gradient of the pressure distribution. This means that the deflection angles caused by diagnostic ultrasound of lower pressure and shorter wavelength could have a range of deflection angles similar to that caused by therapeutic ultrasound. These considerations indicate the possibility of using the technique to visualize both diagnostic and therapeutic ultrasound fields. For diagnostic applications, characterization of a pulsed ultrasound field is important because it directly determines performance of diagnostic ultrasound equipment. For therapeutic applications that use pulsed ultrasound, for example, extracorporeal shock wave therapy (Chung and Wiley 2002) and sonoporation (Kudo et al. 2009), visualization of an individual pulse shape is also important because the shape determines its therapeutic effects.
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Advantages of use of shadowgraphy Elimination of non-deflected light by image subtraction is another advantage of the proposed technique because elimination by spatial filtering requires careful adjustment of the high-quality optical system. Several studies have indicated the usefulness of image subtraction in the Schlieren technique for elimination of the timeindependent background (Azuma et al. 2002; Neumann and Ermert 2006a, 2006b). However, this subtraction has a different meaning from ours, because these techniques use Schlieren images from which nondeflected light is already removed. An inhomogeneous intensity distribution of illumination light causes artifacts in visualized images. Normalization of a subtracted image by one of the images used for the subtraction is effective (Omura et al. 2011; Shimazaki et al. 2012) if intensity fluctuation inside the light deflection area is negligible. However, we did not carry out normalization because our illuminating light beam had a significant inhomogeneity in intensity caused by interference of coherent light emitted from the laser diode. Advantages of use of focused shadowgraphy Use of a macro lens with a smaller DOF is also key. This enables imaging using large deflection light, which is hardly used in the Schlieren technique, and achieves acquisition of focused shadowgrams that vary with the position of the virtual screen. A flexible setting of the screen position is important because the presence of a physical screen disturbs the acoustic fields beside the screen. As illustrated in the results (Figs. 6 and 7), screen position affects both the sensitivity and distortion of visualized field images, and there is a trade-off between them (Fig. 8). This means that ultrasound fields of a very wide pressure range from IVUS to high-intensity therapeutic ultrasound can be visualized using one optical system simply by changing the virtual screen position. High-frequency ultrasound fields can also be visualized by simply increasing the magnification of the focusing optics (Fig. 9, Supplemental Video 1). The pixel size of the images taken using the macro lens placed at a WD of 0.25 m was 8 3 8 mm2. The simple optics enables combination with optical microscopy. For example, the pixel size of 0.93 3 0.93 mm2 was achieved using a 310 objective lens (Kudo et al. 2010). Limitations Image distortion caused by light deflection is the major limitation. This means that the proposed technique cannot replace hydrophone measurements. Shadowgraphy visualizes the second spatial derivative of an inhomogeneous refractive index distribution (Settles 2001),
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and the distribution can be reconstructed by double integration, in principle. Yet, it is difficult in practice because a shadowgram essentially contains image distortion caused by light deflection. Use of focused shadowgraphy reduces the distortion by placing the virtual screen closer to the ultrasound focus; however, reconstruction is still difficult because a shadow projected on a physical screen (Miyasaka et al. 2014) is different from an image captured using the camera with a finite DOF. Further studies are ongoing on this point. Under the condition of higher ultrasound intensity, non-linear steepening causes a rapid increase in light deflection angle. Light rays with larger deflection angles are eliminated by the pupil of the camera lens, and information on the steep pressure change is lost from the visualized images. Non-linear steepening can be observed in the waveform in Figure 4 (p1/2 5 17.8 MPa, 21.3 MPa), and the spatial peak temporal average intensity in the pressure cycle containing the peak positive and negative pressures was 174 W/cm2. Further increase in ultrasound intensity leads to cavitation phenomena, and light scattering caused by the cavitation bubbles disturbs field visualization. Pitts et al. (1994) proposed techniques that can determine an ultrasound field more precisely by detecting a phase shift of light integrated over its path transmitted through an ultrasound field. Two techniques, the phase recovery technique (Pitts et al. 1994) and phase contrast technique (Harigane et al. 2013; Pitts et al. 2001), were used for phase detection. The phase recovery technique estimates the phase shift by an iterative calculation using a magnitude image acquired at a known distance. The phase contrast technique derives the phase shift by using a magnitude image acquired using a p/2 delay column placed at the position of the zeroth-order light. These techniques can reconstruct 3-D fields using a CT algorithm in the pressure range of tens to hundreds of kilopascals in which light deflection is negligible. However, use of largely deflected light in shadowgraphy makes 3-D reconstruction difficult. CONCLUSIONS A simple technique suitable for visualization of medical ultrasound fields was proposed. The principle underlying the technique was described in comparison with two conventional techniques, the Schlieren technique and shadowgraphy, and the advantages and disadvantages were discussed. Two types of acoustic fields, a focused field and a high-frequency field, of short-pulsed ultrasound were visualized to confirm the practical usefulness of the proposed technique. The experimental results indicated that alternating pressure change in traveling short-pulsed ultrasound fields can be visualized
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clearly. Sensitivity and spatial resolution can be controlled flexibly simply by changing the focus position and magnification of the camera lens, indicating the usefulness of the technique for visualization of various medical ultrasound fields from small high-frequency IVUS ultrasound fields to high-intensity therapeutic ultrasound fields. The new shadowgraph technique cannot replace hydrophone measurements; however, this simple technique is useful not only for engineering laboratories, but also for the teaching environment to help laboratory staff who are not familiar with ultrasound understand its nature. Acknowledgments—The author expresses sincere gratitude to the late Professor Katsuyuki Yamamoto for his scrupulous discussion that initiated the research. Supportive comments from Professor Koichi Shimizu were invaluable. The author is also grateful to Professor Shin-ichiro Umemura for permission to use his Schlieren system. This study was supported by JSPS KAKENHI Grant 23300182.
SUPPLEMENTARY DATA Supplementary data related to this article can be found at http:// dx.doi.org/10.1016/j.ultrasmedbio.2015.03.004
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