Ординатура / Офтальмология / Английские материалы / Ultrasonography of the Eye and Orbit 2nd edition_Coleman, Silverman, Lizzi_2006

Figure 2.23. High-frequency M-mode images of an iris vessel in a rabbit's eye, taken over about four cardiac cycles. Left: In grayscale image, stationary structures remain at constant range, whereas flowing blood particles change in range with time. Right: Colorized image of same data illustrates pulsatile flow, with some regurgitation during diastole. (see color image)

M-MODE

M-mode (Figure 2.23) represents a cross between A- and B-modes. As in A-mode, the transducer interrogates a single line of sight, but as in B-mode, a two-dimensional image is formed. In M-mode, however, the vertical axis represents time rather than lateral position (as it does in B-mode). M-mode is useful for demonstration of tissue motion. Stationary tissue structures will maintain a constant range from the transducer, so echoes will appear vertical on the screen. Where tissue motion occurs (i.e., vessel wall motion) the range will vary with time, and this will be evident in the M-mode image. M-mode is not generally available on commercial ophthalmic systems.

SWEPT-MODE

Swept-mode combines M-mode and B-mode (Figure 2.24) (23,24). In conventional B-mode imaging, vectors are usually placed a beam width apart or less. In swept-mode, vectors are placed much less than a beam width apart. The advantage of this is that groups of adjacent vectors within a beam width of each other can be treated as viewing the identical spatial position over time (time being related to the pulse repetition frequency), while vectors that are more than a beam width apart allow

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formation of a conventional B-mode image. Thus, a swept-mode image is equivalent to a B-mode image composed of overlapping M-scans. Swept-mode is thus capable of showing tissue motions, including blood flow, in the context of a B-mode image. The advantage of swept-mode in comparison to Doppler is that flow information is obtained at the same high resolution as the underlying B-mode data. Disadvantages include lower sensitivity and slower frame rate.

Figure 2.24 Swept-mode is, essentially, a B-mode image in which vectors are spaced much closer together than the beam width. Because of this, groups of adjacent vectors within a beam width of each other are not spatially independent. This allows treating these groups of overlapping vectors as local M-scans. The uppermost 50-MHz image of the angle region in a rabbit's eye was constructed from 128 vectors spaced 18 microns (about 4 vectors/beam width). This has the appearance of a conventional B-mode image. The center image is of the same tissue but with 1,024 vectors spaced 2.2 microns apart (about 30 vectors/beam width). In this highly oversampled case, we can see areas where the echo phase is decorrelated compared to surrounding tissues (arrows). This effect results from blood flow, where blood cells change in range over time. Based on the PRF and the change in range per vector, flow velocity can be computed. The bottommost figure is a color-flow image generated from the digitized echo data. (see color image)

LINEAR ARRAY SYSTEMS

At the time of this writing, ophthalmic ultrasonography is almost universally performed using mechanical sector scan probes, a technology that has almost disappeared outside of this specialty. General purpose ultrasound instruments rely on linear array transducers for B-mode imaging. This distinction is attributable to a number of factors. Ophthalmic ultrasound is performed at frequencies that are generally higher than those used in other specialties, with a few exceptions. Fabrication of arrays and control circuitry become more difficult and expensive as frequency increases. These factors, taken in the context of the relatively small ophthalmic ultrasound market, have kept ophthalmic ultrasound out of the technologic mainstream.

Small-parts linear array probes with a center frequency of 10 MHz or more are available. One should consider the advantages of linear array technology. These include:

High frame rate

Large effective aperture—improves lateral resolution

Dynamic focusing—movable/multiple synthetic focal zones

Special scan modes

Continuous wave Doppler

Color flow Doppler

Power Doppler

Tissue harmonic imaging

The trend toward decreasing costs, greater compactness, and increasingly higher available frequencies for linear array systems suggests that this technology is likely to have an increasing impact on the performance of ophthalmic ultrasonography in the future.

OTHER SCAN MODES

DOPPLER MODES

The Doppler effect (25) is well known: If a sound source is moving toward the listener, the wavelength is compressed, and the pitch increased (Figure 2.25). The

opposite effect occurs when the sound source moves away from the listener. This effect has been used in ultrasound systems for measurement and visualization of blood flow and in ophthalmology for visualization and quantification of flow in the orbital vessels and tumors (26). The Doppler frequency shift, fd, is, by definition,

the difference between the emitted (fe) and received (fr)

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frequencies, that is, fd = fe-fr. It is determined by fd = fe[2v ÷ (c-v)], where c represents the speed of sound, and v represents the component of the velocity of the

scatterer (e.g., blood cells) along the transducer beam axis. For instance, if a vessel has a flow velocity of 10 cm sec-1, then the Doppler frequency shift is 10 ×

106[0.02 ÷ (1540-0.01)] = 130 Hz for a 10-MHz source. Notice that this frequency is in the audio range. Doppler systems provide an audio output that allows the sonographer to “hear” flow.

Figure 2.25. The wavelength of reflections from a particle in motion toward the transducer is shortened.

In Doppler ultrasonography, our interest is detection of frequency shifts associated with tissue motion. As such, Doppler becomes more sensitive as transducer bandwidth is reduced or, equivalently, as pulse duration is increased. This means that as Doppler resolution increases, spatial resolution decreases.

The most basic Doppler mode is “continuous wave” (CW) Doppler. In CW Doppler, the pulser is replaced by an oscillator that produces a continuous sine wave voltage that excites the transducer. Because this transducer is exclusively generating a continuous emission, a separate transducer is used to receive echoes. Alternatively, in linear array systems, one subset of elements can be used to emit, while another set acts as the receiver. Color-flow or duplex Doppler (Figure 2.26) involves a simultaneous display of a B-mode image with superimposed color information indicating areas of flow. In color-flow, a subset of elements in a linear array emits pulses several cycles in duration. This allows measurement of Doppler shift as well as range and direction simultaneously (but with lower spatial resolution than the underlying B-mode image and with reduced Doppler sensitivity compared to CW). The color-flow image can be used to select a vessel for CW interrogation of the Doppler waveform, as shown in Figure 2.27.

Doppler electronics (Figure 2.28) differ in some ways from that of A- and B-mode systems. After the echo data are amplified, they are processed by a component called a mixer. The mixer multiplies the amplified echo waveform by the excitation sine waveform. This provides a signal in which the sum (fe + fr) and the

difference (fe-fr) of the frequency components of the two inputs are combined. This signal is then bandpass filtered to remove the summation component, leaving

only fe-fr, which is,

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of course, equivalent to fd, the Doppler frequency shift. It may also be necessary to perform a high-pass filter operation of the Doppler signal to remove extraneous

signal components associated with tissue motion. The combined amplifier and filter components are referred to as a demodulator. Because fd is in the audio range, this signal is then amplified with an audio frequency amplifier.

Figure 2.26. Color-flow Doppler image of a normal human eye. The scan was taken with the transducer in a

vertical orientation such that superior is toward the right. The image shows the central retinal vessels as well as

the ciliary artery just superior to the nerve. (see color image)

Figure 2.27. The CW Doppler waveform associated with a vessel is obtained by choosing the area of interest

on a static color-flow image and adjusting the measurement angle (to perform cosine correction) where

necessary. (see color image)

Figure 2.28. Schematic drawing representing the electronic components used to generate directional color-flow

information.

The previously shown arrangement, however, provides only the magnitude of the Doppler frequency shift. Because we are often interested in the direction of flow as well, a somewhat more complicated electronic processing method called phase quadrature detection is used. In this method, the echo signal is passed to two separate demodulators, the first of which (as before) multiplies the signal by the excitation waveform, and the second multiplies the echo by an excitation waveform

that has been rendered 90 degrees out of phase. These are referred to as the in-phase and the quadrature signals, respectively. The in-phase and quadrature channels allow determination of whether fd is positive or negative; if fd is positive, the quadrature channel lags behind the in-phase channel by a phase shift of 90

degrees, whereas, if fd is negative, then the quadrature channel is 90 degrees advanced, in respect to the in-phase channel.

CW Doppler systems provide a graphic representation of flow by converting the Doppler signals into positive and negative velocity values. These are plotted over a period of several cardiovascular cycles to provide a good representation of the systolic and diastolic blood flow pattern in a vessel.

In color-flow Doppler systems, colored pixels representing flow are superimposed onto the B-mode image. A color-scale, usually ranging from reds (representing arterial flow, i.e., flow toward the transducer) to blues (venous flow) is presented on the display. The sonographer can adjust several parameters to optimize the color-flow presentation, including the range of velocities to be displayed, the write priority of color flow versus gray-scale information, and filtering functions (Wall filters) used to suppress Doppler shifts associated with motions of solid tissues, such as vessel walls or respiratory motions. In addition, the user can generally modify the PRF used for acquiring the Doppler signal. This is significant in that the highest Doppler frequency that can be accurately characterized is one half of the PRF. If this is exceeded, then a phenomenon called aliasing occurs.

Color flow is advantageous because vessels (in the scan plane) are seen in the context of the B-mode image, which facilitates identification. Also, color-flow

imaging allows estimation of the angle of the vessel in relationship to the acoustic beam axis. Because Doppler systems can provide only a measurement of the velocity component of flow in the beam axis, a cosine adjustment term must be used to correct Doppler velocity values, vd = v cos(?), where ? is the angle between

the transducer axis and the flow direction and, vd is the uncorrected Doppler velocity value. Color-flow Doppler systems allow the user to indicate vessel orientation

so that the cosine correction term can be applied to the uncorrected Doppler velocity values.

Power Doppler (27) is also provided on most instruments to perform color-flow Doppler. In Power Doppler, the Doppler frequency shift signal is integrated. This has the effect of removing directional and velocity information but provides a color-flow map of perfusion that has a significantly higher sensitivity than conventional color-flow Doppler and less sensitivity to angular orientation. Power Doppler is particularly useful in a situation of slow-flow and tortuous vasculature, as in some tumors.

TISSUE HARMONIC IMAGING

Tissue harmonic imaging (THI) was developed in the late 1990s (28). It was discovered serendipitously as a consequence of attempts to develop a means for improved detection of flow, using ultrasound contrast agents. Such agents consist of microbubbles, lipid shells filled with air, or other substances with a high acoustic impedance inhomogeneity compared to blood. It was anticipated that the microspheres would resonate at specific ultrasound frequencies, and that this would result in emission of echo data from the microspheres at harmonics of the emitted ultrasound frequency. By filtering out the fundamental emitted frequency, then, tissue echoes would be suppressed, while the harmonic vibration modes of the microspheres would be detectable. Although this process was found to be valid, early users discovered that even in the absence of contrast agent, tissues were seen with better contrast and clearer borders in images made at the second harmonic (i.e., double the emission frequency). This effect is a result of nonlinear interaction between the ultrasound pulse and the tissues through which it propagates. The speed of sound is affected by the density of the material through which it propagates. A sound wave, however, is by definition a pressure wave, with compressive and decompressive components. Thus, as a sound wave travels through a medium, the compressed component of the

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pulse tends to move slightly slower than the decompressed part of the pulse. This distortion is a function of both the tissues through which the pulse travels and the overall distance—the effect is cumulative with range. The distortion of the pulse waveform results in generation of harmonics. The reason that this is interesting is that the harmonic has comparatively small side lobes compared to those of the fundamental. Because side lobes effectively reduce lateral resolution, generating an image at the harmonic produces more well-defined boundaries.

Tissue harmonic imaging requires a transducer with sufficient bandwidth to capture at least the second harmonic. This is generally accomplished by exciting the transducer at two thirds of its center frequency and receiving at four thirds of center frequency. The simplest technology for producing tissue harmonic images is by using bandpass filters centered at the second harmonic that filter out the fundamental. The disadvantage of this approach is that both bands must have fairly narrow bandwidths, and hence poor axial resolution, to achieve this separation. An alternative approach is pulse-inversion. In this method, two pulses are emitted

in quick succession, one of normal phase and the other inverted. When echoes from the two pulses are acquired and added, the fundamental and all odd harmonics are eliminated, leaving only the even harmonics. This provides an effective means of capturing and generating images at the second harmonic.

THI is now widely incorporated into linear array imaging systems. In fact, it is so effective that it is often used as the default imaging mode.

SPECTRAL PARAMETER IMAGING

The interaction between an ultrasound pulse and the tissue through which it propagates causes the reflected or backscattered signal to differ from the emitted signal. An extensive literature exists regarding the effect of tissue microarchitecture on backscatter (29, 30, 31, 32, 33). It is understood, for instance, that as tissue inhomogeneities become smaller compared to a wavelength, they more effectively scatter the higher frequencies present in the ultrasound pulse (Figure 2.29). (This is the same physical principle that causes the sky to be blue.) Inhomogeneities that are much smaller than a wavelength become Raleigh scatterers, with backscatter increasing with the fourth power of frequency. Also known, and intuitively obvious, is that as the number of scatterers per unit volume increases, backscatter increases as well. However, if the scatterer concentration rises high enough, the scatterers effectively become background instead of foreground, and the spaces between them behave as the scatterers. The geometric form of scatterers (spherical, filamentous, lamellar) also affects the backscattered signal. In the case of nonisotropic scatterers (filaments or lamellas), their orientation relative to the ultrasound beam is an important consideration. Clearly, this is a complex process. Nevertheless, if certain simplifying assumptions are made (e.g., weak scatterers and negligible attenuation), quantitative estimates of scatterer size and concentration can be derived from measurement of the difference between the power spectrum of the emitted pulse and the received echo, referred to as the calibrated power spectrum (CPS). Because the CPS is typically quasilinear in appearance, the linear best fit to the CPS is used as a means for characterizing tissues.

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Mathematical modeling of acoustic backscatter shows that the slope (dB/MHz) of the linear best fit to the CPS is directly related to scatterer size, whereas the intercept (dB) relates to scatterer concentration and relative impedance. By measuring spectra at successive spatial positions within a B-mode image, estimates of these quantities can be made and the gray-scale pixel values replaced by colors representing calculated scatterer size and concentration. This approach is used in ophthalmology to characterize tissues, such as tumors, hemorrhage, and corneal scars. Spectral parameter images (Figure 2.30) provide a visual representation of the physical properties of tissues and also quantitative values related to the mean value and variation of each parameter that allow comparison with cases of known pathology or determination of changes occurring over successive examinations.

Figure 2.29. Mathematical modeling of acoustic backscatter has shown that as scatterer size increases from much less than a wavelength to a half-wavelength, spectral slope (reflected amplitude versus frequency) goes progressively from positive to negative values.

Figure 2.30. Spectral parameter images generated by determining calibrated spectra along each vector and replacing the grayscale pixel values representing the envelope of the echo data with color values representing a spectral parameter. In this case, three kinds of spectral parameter images were generated from a high-frequency arc-scan of the anterior segment of the eye with hyphema. Upper left: A midband fit image is presented in grayscale in the correct geometric format. The other three images show images in stretched rectilinear format. Bottom left: Midband fit. Upper right: Slope. Lower right: Slope. Lower right: Intercept.

Midband fit allows reduction in speckle and other source of noise.

REFERENCES

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2.Baum G, Greenwood I. The application of ultrasonic locating techniques to ophthalmology: part 2. Ultrasonic visualization of soft tissues. Arch Ophthalmol 1958;60:263-279.

3.Ossoinig KC. Quantitative echography: the basis of tissue differentiation. J Clin Ultrasound 1974;2:33-46.

4.Sanders DR, Kraff MC. Improvement of intraocular lens power calculation using empirical data. J Am Intraocul Implant Soc 1980;6:263-267.

5.Binkhorst RD. Intraocular lens power calculation. Int Ophthalmol Clin 1979;19:237-252.

6.Holladay JT, Prager TC, Ruiz RS, et al. Improving the predictability of intraocular lens power calculations. Arch Ophthalmol 1986;104:539-541.

7.Hoffer KJ. Pre-operative cataract evaluation: intraocular lens power calculation. Int Ophthalmol Clin 1982;22:37-75.

8.Aslanides IM, Aslanides MN, Reinstein DZ, et al. Have you ever seen a pachyderm [Letter]? J Refract Surg 1995; 11:162-164.

9.Flanagan G, Binder PS. Estimating residual stromal thickness before and after laser in situ keratomileusis. J Cataract Refract Surg 2003;29:1674-1683.

10.Ossoinig KC. Standardized echography: basic principles, clinical applications, and results. Int Ophthalmol Clin 1979;19:127-210.

11.Zaldivar R, Shultz MC, Davidorf JM, et al. Intraocular lens power calculations in patients with extreme myopia. J Cataract Refract Surg 2000;26:668-674.

12.Silverman RH, Reinstein DZ, Raevsky T, et al. Improved system for ultrasonic imaging and biometry. J Ultrasound Med 1997;16:117-124.

13.Reinstein DZ, Silverman RH, Raevsky T, et al. Arc-scanning very high-frequency ultrasound for 3-D pachymetric mapping of the corneal epithelium and stroma in laser in situ keratomileusis. J Refract Surg 2000;16:414-430.

14.Coleman DJ, Silverman RH, Chabi A, et al. High resolution ultrasonic imaging of the posterior segment. Ophthalmology, 2004;111:1344-1357.

15.Rosenfeld A, Kak AC. Digital Picture Processing. New York: Academic Press; 1982.

16.Cusumano A, Coleman DJ, Silverman RH, et al. Three dimensional ultrasound imaging: clinical applications. Ophthalmology 1998;105:300-306.

17.Silverman RH, Coleman DJ, Rondeau MJ, et al. Measurements of ocular tumor volumes from serial, cross-sectional ultrasound scans. Retina 1993;13:69-74.

18.Romero JM, Finger PT, Rosen RB, et al. Three-dimensional ultrasound for the measurement of choroidal melanomas. Arch Ophthalmol

2001;119:1275-1282.

19.Delcker A, Martin T, Tegeler C. Magnetic sensor data acquisition for three-dimensional ultrasound of the orbit. Eye 1998;12:725-728.

20.Pavlin CJ, Sherar MD, Foster FS. Subsurface ultrasound microscopic imaging of the intact eye. Ophthalmology 1990;97:244-250.

21.Pavlin CJ, Harasiewicz K, Sherar MD, et al. Clinical use of ultrasound biomicroscopy. Ophthalmology 1991;98: 287-295.

22.Pavlin CJ, Harasiewicz K, Foster FS. Ultrasound biomicroscopy of anterior segment structures in normal and glaucomatous eyes. Am J Ophthalmol 1992;113:381-389.

23.Kruse D, Fornaris J, Silverman R, et al. A swept-scanning mode for estimation of blood velocity in the microvasculature [Letter]. IEEE Trans Ultrason Ferroelectr Freq Control 1998;45:1437-1440.

24.Silverman RH, Kruse D, Coleman DJ, et al. High-resolution ultrasonic imaging of blood-flow in the anterior segment of the eye. Invest Ophthalmol Vis Sci 1999;40: 1373-1381.

25.Wells PN. Ultrasonic colour flow imaging. Phys Med Biol 1994;39:2113-2145.

26.Tanquart F, Berges O, Koskas P, et al. Color Doppler imaging of orbital vessels: personal experience and literature review. J Clin Ultrasound 2003;31:258-273.

27.Macsweeney JE, Cosgrove DO, Arenson J. Colour Doppler energy (power) mode ultrasound. Clin Radiol 1996;51:387-390.

28.Duck FA. Nonlinear acoustics in diagnostic ultrasound. Ultrasound Med Biol 2002;28:1-18.

29.Lizzi FL, Greenebaum M, Feleppa EJ, et al. Theoretical framework for spectrum analysis in ultrasonic tissue characterization. J Acoust Soc Am 1983;73:1366-1373.

30.Insana MF. Ultrasonic imaging of microscopic structures in living organs. Int Rev Exp Pathol 1996;36:73-92.

31.Hosokawa T, Sigel B, Machi J, et al. Experimental assessment of spectrum analysis of ultrasonic echoes as a method for estimating scatterer properties.

Ultrasound Med Biol 1994;20:463-470.

32.Hunt JW, Worthington AE, Kerr AT. The subtleties of ultrasound images of an ensemble of cells: simulation from regular and more random distributions of scatterers. Ultrasound Med Biol 1995;21:329-341.

33.Lizzi FL, Astor M, Feleppa EJ, et al. Statistical framework for ultrasonic spectral parameter imaging. Ultrasound Med Biol 1997;23:1371-1382.

Authors: Coleman, D. Jackson; Silverman, Ronald H.; Lizzi, Frederic L.; Lloyd, Harriet; Rondeau, Mark J.; Reinstein, Dan Z.; Daly, Suzanne W. Title: Ultrasonography of the Eye and Orbit, 2nd Edition

Copyright ©2006 Lippincott Williams & Wilkins

> Table of Contents > 3 - Ocular Diagnosis

3

Ocular Diagnosis

HISTORICAL BACKGROUND

The first use of ultrasound for ophthalmic diagnosis was reported in 1956 by Mundt and Hughes (1), who used industrial ultrasound equipment to examine enucleated normal eyes and eyes with intraocular tumors. The first clinical use of A-scan in ocular diagnostic problems was described in 1957 by Oksala (2), in the first of many pioneering papers. Later, Jansson (3) of Sweden described the use of ultrasound for ocular measurement and made critical in-vitro measurements of the sound velocity constants for ocular tissues (4). Oksala (5) also made early measurements of sound velocities of tissues. Sorsby (6), using A-scan, compared axial lengths in a large patient population and described sex and age differences. Coleman and Carlin (7) described the first axial length measurements of the eye, using an electronic interval counter to make precise axial measurements and to document lens movement in accommodation. Giglio (8) also described axial measurements, using a similar system.

Oksala's initial work was followed by many subsequent papers on A-scan for clinical diagnosis (9). Later, Bronson (10) developed ultrasonically directed intraocular forceps for foreign bodies. Ossoinig (11) popularized a specific form of A-scan equipment and described clinical results with many original observations on the A-scan properties of specific tissues. He developed a sophisticated diagnostic technique that emphasized the quantification of echo amplitudes, using a tissue standard and “s” shaped amplification of the A-scan, which he termed standardized echography. He also described kinetic A-scans in which movement of both the transducer and vascular structures is used to characterize tissues (12). This technique remains in widespread use.

Other investigators who have contributed early, original observations to A-scan diagnosis include Buschmann (13) and Gernet (14) of West Germany, Massin and Poujol (15) of France, Francois and Goes (16) of Belgium, Vanysek and Preisova (17) of Czechoslovakia, Bertenyi (18) of Hungary, and Gallenga (19) of Italy. In the United States, in the hospital-based laboratory at the Wills Eye Institute, Sarin et al. (20), under the direction of Keeney (21), made early contributions regarding A-scan evaluation. At the Walter Reed Army Hospital, Penner and Passmore (22) and Cowden and Runyon (23) described the uses of A-scan in the diagnosis of foreign bodies. Coleman (7) demonstrated high frequency (25-MHz A-scan) evaluation of the choroid to measure the in-vivo thickness of this highly vascular erectile tissue.

B-scan diagnosis was first developed by Baum and Greenwood (24, 25, 26) in 1958. They made numerous original observations on B-scan evaluation of the eye and orbit. Their work, featured on the cover of the Journal of the Acoustical Society of America (27) of a B-scan of the eye, was the adumbration of developments to come. Baum's efforts were devoted primarily toward the development of equipment with increased accuracy and better resolution (Figure 3.1).

Purnell and Sokollu (28) used a similar prototype B-scan (developed by General Precision Instruments) and made many seminal observations that had a major influence on B-scan diagnosis (Figure 3.2). His laboratory described orbital B-scan evaluation and provided the first systematic classification of orbital disease with B-scan ultrasonography (29). He and others were the first to use the magnetic properties of a foreign body in ultrasonic diagnosis (28). Purnell and Sokollu (30,31) also described special techniques for using continuous wave ultrasound (generally related to therapeutic applications) in early experiments for the treatment of retinal detachment with ultrasound. They developed the first handheld contact B-scanner for use in ophthalmic ultrasound diagnosis (which preceded the Bronson contact

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ultrasound instrument), but it was never marketed (Figure 3.3) (32).