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

maximum value; this width is approximately equal to 0.5 (?/a)R. Often it is more meaningful to specify the angle that describes the 3-dB periphery of the main lobe (Figure 1.25); this angle (in radians) is approximately equal to 0.5 (?/a).

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Figure 1.23. Relative pressure amplitude as function of distance along tranducer's central axis. Amplitude

decreases monotonically for distances larger than a2/?.

Figure 1.24. Dependence of far-field pressure amplitude upon angular position. Plot is drawn for points at a fixed distance, R, from transducer center; abscissa values are proportional to off-axis distance, r.

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Figure 1.25. Beam pattern of an unfocused transducer showing near-field and far-field regions. The 3-dB angular width of the main lobe is 2 ?.

Ultrasonic wavelength and transducer dimensions determine both near-field and far-field characteristics. As ? decreases (frequency increases), the length of the near field increases and the far-field beamwidth becomes narrower. As an example, if a is 2.5 mm and ? is 0.1 mm (15 MHz), the near-field length is 62.5 mm, and the far-field beam angle is 1.2 degrees.

Ocular examinations with unfocused transducers are usually performed within the near field, where the beam is typically several mm wide. This degree of lateral

resolution is too coarse for many ophthalmic examinations, especially B-scans, can be improved by focusing, as discussed later.

Focused Transducers

Transducers are focused by using acoustic lenses, as discussed in connection with refraction. However, the width of the resultant beam is not zero at the focal point, as diagrammed in Figure 1.17; rather, diffraction causes a small, but finite, beamwidth that depends on ultrasonic wavelength and transducer dimensions.

Figure 1.26. Beam pattern of focused transducer. Dotted line indicates focal plane.

Classic analyses treating Huygens' point sources show that focused beams have profiles, such as those shown in Figure 1.26 (30). In the focal plane, the beam exhibits a lobed structure of the same type encountered under unfocused far-field conditions. Here, the side lobes straddle the narrow-focused region, and the main lobe beamwidth is equal to 0.5(?/a)F, where F is the focal length of the transducer and a, again, represents the radius of the transducer rim. Typically, ? is 0.15 mm (10 MHz), a is 5 mm, and F is 30 mm, resulting in a beamwidth of 0.45 mm. The same transducer dimensions provide smaller beamwidths as frequency is increased.

In the selection of clinical focused transducers, the desired beamwidth can be obtained by proper selection of wavelength and transducer parameters (a and F). However, several other factors must be considered when choosing a transducer that will permit adequate tissue examination. Focal lengths must be chosen to provide focusing in the desired tissue region; focal lengths near 30 mm allow for intraocular focusing, whereas larger focal

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lengths are needed for deeper orbital examinations. Furthermore, high frequencies cannot be used in the orbit because of attenuation. These combined limitations on focal lengths and frequencies constrain lateral resolution in the orbit more severely than in the eye. In addition, clinical transducers must be weakly focused to generate narrow beams over relatively long tissue depths. Strong focusing, obtained with large transducer diameters, is not usually desirable; although strong focusing produces narrow beamwidths, it can do so only over unacceptably shallow tissue depths. The specification of an f-number (equal to F/2a) summarizes these considerations for ultrasound in the same manner used in optics. Small f-numbers (e.g., 1.5) provide sharp focusing over limited depths; large f-numbers (e.g., 3-4) provide moderate focusing over longer depths.

The same considerations apply to very high frequency (VHF) transducers used in 40-MHz examinations of anterior chamber structures. For example, transducers with 3-mm radii and 12-mm focal lengths can produce beamwidths near 50 µm for very fine lateral resolution. However, their depths of focus can be less than 1 mm, so that care must be used in positioning the transducer focal zone over the tissue segment to be examined.

These discussions of unfocused and focused transducer beam patterns are useful in all cases, but they are rigorously applicable only when the ultrasonic pulse contains several cycles of oscillation. They must be modified, if the transducer generates a very short pulse, such as the single-cycle pulse illustrated in Figure 1.20. Short pulse durations do not allow enough time for standard interference patterns to develop. For a single-cycle pulse, the 3-dB main lobe width is equal to that quoted previously, but side lobe patterns and near-field characteristics differ, in several respects, from those discussed previously (31).

COMPOSITE RESOLUTION

Composite (axial and lateral) resolution depends on frequency and transducer geometry. These factors, in turn, are influenced by the type of examination to be made. As a brief summary of preceding sections, anterior segment examinations present an environment where fine resolution can be achieved; ocular examinations permit fine resolution; orbital examinations must be carried out at lower resolution.

For an ideal transducer, both components of resolution improve as resonant frequency is increased; however, attenuation also increases with frequency, limiting the values that can be realized in practice. In the anterior segment, frequencies near 40 MHz have provided axial resolution of 30 µm and lateral resolution near 50 µm. Within the eye, frequencies near 20 MHz have been used to attain an axial resolution of 0.1 mm; focal lengths of 30 mm, together with a transducer radius of 5 mm, allow lateral resolution of 0.2 mm. In orbital examinations, increased absorption currently limits frequencies to a maximum near 10 MHz. These lower frequencies and necessarily long focal lengths (e.g., 60 mm) reduce axial resolution to approximately 0.3 mm and lateral resolution to 0.9 mm. Newer, more sensitive transducers promise to alleviate some of these restrictions, especially in anterior regions of the orbit (32).

Transducer Arrays

Single-element transducers, as shown in Figure 1.1, are a standard use in ophthalmology, and mechanical scanning is used to generate B-mode images. Other medical specialties frequently use piezoelectric arrays for electronic focusing and scanning (4,33). Arrays comprise a set of small, discrete elements that transmit and receive ultrasonic pulses, emulating Huygens' sources that can be individually controlled to electronically focus and steer the overall beam. Electronic beam control provides rapid, versatile operation, without the need for mechanical scanning. However, arrays increase the complexity and cost of ultrasonic instruments. For high-quality performance, the sizes and spacings of array elements must be comparable to or smaller than the ultrasonic wavelength; arrays are common in systems using frequencies of, for example, 7 MHz, but they have not yet found widespread use at the higher frequencies and smaller wavelengths used in ocular examinations.

There are three basic types of arrays: linear arrays (for linear scanning), phased arrays (for sector scanning), and annular arrays (for controlling focal zones along single scan lines). Each of these controls the timing of excitation pulses to focus the transmitted pulse and also applies time delays to returned echoes to focus received signals.

A linear array comprises a set of thin, parallel, rectangular elements on a planar substrate. On transmit, focusing is achieved by exciting each element in a programmed sequence. This is shown schematically for three array elements in Figure 1.27. Elements at the edge of the desired beam are excited first, and the

central element is excited last. The timing of the excitations is adjusted so that the pulses from all elements arrive simultaneously and in phase at the desired focal point, producing a large focal-point pressure pulse (at time t2 in the figure). In the receive mode, echo signals from the elements are time-shifted in a similar

manner; the maximum time delay is applied to the central element, so that echoes from the focal point are aligned in time. The shifted RF echoes are then summed to obtain the desired receive focusing. Arrays permit dynamic focusing, which extends the effective depth of focus on receive. In this mode, the applied echo time-shifts are continually adjusted to maintain an effective focus at the distance from which echoes are returning at that time. Accordingly, the effective focal

where the superscript bar denotes an average over time. I is usually specified in watts/cm2. In practice, a set of subscripts denotes the spatial and temporal averaging intervals used in intensity specifications, as subsequently described for the U.S. Food and Drug Administration (FDA) exposure indices.

At high exposure levels, ultrasound can alter tissues by thermal effects, mechanical phenomena, and cavitation (34). Thermal effects arise from absorbed ultrasonic energy, which is converted to heat. When the incident energy is sufficiently high, the corresponding temperature rise can damage or denature tissue constituents. Mechanical effects can occur when the incident beam is absorbed or reflected by a tissue structure. These phenomena redirect the beam's momentum and generate radiation forces that can produce tissue motion or fluid streaming at high intensities. Cavitation, which occurs most readily at low frequencies, can occur with large ultrasonic pressure oscillations; in these cases, the negative pressure may promote the formation of gas-filled microbubbles that can grow until a positive pressure cycle causes their sudden collapse. The collapse can be accompanied by large, mechanical forces locally disrupting tissues in the vicinity of the bubble.

In view of the widespread application of diagnostic ultrasound, the low intensities and pressures used in current systems pose no known threat of tissue damage. Animal studies, the absence of reports of clinical damage, and extrapolation of laboratory data all point to the safety afforded by present diagnostic systems. FDA guidelines, described in the following section, have been formulated to continue this record of safety. At high exposure

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levels, however, animal studies reveal that physical alterations can be produced in ocular tissues.

Effects of high-intensity ultrasound on ocular tissues have been studied since 1938, when Zeiss (31) described in-vitro cataract production. A series of investigations has shown that thermal phenomena can produce cataracts (35, 36, 37, 38) and lesions of the cornea (39), choroid, retina, and sclera (40, 41, 42). Mechanical vascular compression can potentiate chorioretinal lesion production by suppressing blood-flow cooling (42). The exposure levels needed to produce ocular damage were orders of magnitude larger than those used in diagnostic systems. This finding is consistent with a report that found no ocular damage after diagnostic exposures as long as 4 hours (43). Animal studies have also found no damage following prolonged, elevated exposures using VHF (40 MHz) diagnostic systems(44).

High-intensity focused ultrasound (HIFU) has been studied as a means of exploiting these physical interactions to treat diseases in various organs (45,46). Typically, HIFU treatments use a series of focal lesions produced, using exposures of several hundred watts/cm2, with durations of several seconds; these exposures are designed to produce desired effects before blood-flow cooling becomes significant. In the eye, glaucoma has been treated in humans, using thermal ciliary body lesions to decrease aqueous humor production and potentiate alternative outflow pathways (47). In animals, chorioretinal lesions produced retinal adhesion similar to that found using lasers; HIFU exposures prevented the spread of retinal tears and facilitated reattachment of detached retinas (48). In rabbits, vitreous hemorrhages (49) and membranes (50) were successfully disrupted by pulsed HIFU beams, which cause mechanical agitation to promote intravitreal dispersion. Tumor therapy has been investigated, using HIFU to treat human melanoma explants in nude athymic mice (51,52). Ultrasonic hyperthermia has also been applied to tumors, using broad beams with lower intensities (several watts/cm2) to achieve sustained heating (near 45°C) for 30 minutes, for example (53,54).

FDA EXPOSURE INDICES

The FDA has devised safety guidelines to assure that all diagnostic ultrasound devices produce exposure levels that are below specific exposure thresholds (55,56). These indices are defined as follows:

Intensities are specified in W/cm2, using subscripts to denote spatial and temporal factors. The SP (spatial peak) subscript indicates the maximum intensity level in the beam. The PA (pulse average) subscript indicates a temporal average over the duration of a single pulse. The TA (time average) subscript indicates a temporal average over the time interval from one pulse to the next; this interval is set by the pulse repetition frequency (A-mode) or scan rate (B-mode). The subindex 0.3 indicates that the value is “derated” for the effect of attenuation (assumed to have a value of 0.3 dB cm-1 MHz-1) between the transducer and the measurement point.

The mechanical index MI is defined as the derated peak rarefaction pressure, Pr (in megapascals), divided by the square root of center frequency, fc, (in

megahertz). MI is a unitless number related to the risk of cavitation. The thermal index TI is defined as the output power W (in milliwatts) times the center frequency f (in megahertz) divided by 210 mW MHz. The denominator is considered to be the power level required to raise tissue temperature 1°C. Thus, theTI is a unitless number, which at a value of unity indicates that a 1°C temperature increase in the insonified tissue would be expected.

For an ultrasound unit to be sold in the United States, it must meet FDA standards. The FDA provides two tracks under which a diagnostic ultrasound device can

meet these regulatory standards. Under Track 1, the instrument manufacturer demonstrates conformity with 510(k) standards, levels that are deemed to be safe in diagnostic instruments, based on historic experience. For ophthalmology, these levels are as follows: ISPTA.3 = 17 mW/cm2, ISPPA.3 = 28 W/cm2, MI = 0.23 (55). These values are well below those in any other specialty. For instance, the ISPTA.3 thresholds for peripheral vessels, cardiac, and fetal imaging are 720,430, and 94

mW/cm2, respectively, as compared to 17 mW/cm2 in ophthalmology. This conservative ophthalmic threshold is a consequence of the concern for cataract formation, which is relatively high because of the high attenuation coefficient of collagen in combination with the lack of vascular cooling within the lens. Virtually all ophthalmic diagnostic ultrasound units follow Track 1.

Track 3 devices follow the Output Display Standard (56). Under this approach, the instrument must display the MI and/or TI, if conditions exist, under which either might exceed a value of 1.0. For Track 3 ophthalmic systems, the TI must not exceed 1.0, the MI must not exceed 0.23, and the ISPTA.3 must be under 50 mW/cm2.

Track 3 is used most commonly in general purpose instruments that might include a small-parts probe suitable for ophthalmic examinations.

Some debate continues as to the validity of the ophthalmic standards (57), especially for very high frequency (VHF) ultrasound (44). The FDA periodically reviews existing standards in light of current research.

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Figure 1.28. Diagram of schlieren system.

MEASUREMENT OF ULTRASONIC BEAM PARAMETERS

Ultrasonic exposure levels must be determined for some research applications, and manufacturers must specify them to assure compliance with FDA guidelines (58). Small ultrasonic probes (hydrophones) are most often used for accurate measurements of ultrasonic pressure pulses. Needle hydrophones use small transducers as receivers to measure local values of pressure as a function of time. Other hydrophones use large PVDF membranes whose electrodes define small active areas for these measurements; such probes rely on PVDF because it is well matched to water so that their presence does not significantly affect the incident beam. Both types of hydrophone can be calibrated and scanned through the incident beam in a water-filled tank. Their output voltages can be directly related to pressure (in pascals) as a function of time.

Schlieren techniques use an optical system to produce a visible image of an ultrasonic beam. The light intensity at each point in a schlieren image is related to the average pressure amplitude within the imaged beam and provides a semiquantitative measure of beam strength.

In a schlieren system (Figure 1.28), a point source of light and a collimating lens combine to produce a plane wave of light that passes through a fluid-filled optical cell. The light exiting from the cell is focused by an integrating lens upon a small, opaque optical stop. If there are no ultrasonic waves propagating through the fluid in the optical cell, all light is blocked by the stop.

The transducer to be studied is placed in the cell and excited with a continuous-wave sinusoidal voltage. Ultrasonic waves perturb the optical index of refraction within the cell fluid so that the light emerging from the cell is nonplanar and, therefore, is no longer completely focused on the stop. That portion of light bypassing the stop contains spatial and amplitude information relating to beam structure. A reimaging lens converts this information into image form. (If the optical stop were not used, unaffected portions of the incident light would obscure the schlieren image.) Sensitive schlieren systems have been used in a more quantitative fashion to examine pulsed beams.

Radiation pressure techniques have been used to measure the ultrasonic power emanating from a diagnostic transducer. In some implementations, a beam is reflected at 45 degrees by a highly reflective plate. An analytic balance measures the small force on the plate that results from the redirection of wave momentum. This force is directly related to the incident ultrasonic power; a power level of one mW produces a force of 0.067 mg. Carefully designed measuring systems with sensitive balances are capable of measuring the low power levels encountered in diagnostic systems.

In the next chapter, the use of the physical principles in designing and constructing clinical instruments for ocular examination will be discussed.

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49.Lucas BC, Driller J, Iwamoto T, et al. Ultrasonically induced disruption and hemolysis of vitreous hemorrhages. Ultrasound Med Biol. 1989;15:29-37.

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55.Information for Manufacturers Seeking Marketing Clearance of Diagnostic Ultrasound Systems and Transducers. Rockville, MD: Food and Drug Administration, Center for Devices and Radiological Health; 1997.

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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 > 2 - Ultrasonic Systems

2

Ultrasonic Systems

The two most commonly used ultrasonic imaging modalities in ophthalmology are termed A-mode and B-mode. Each presents anatomic information in a distinctive display format. A-mode refers to a graphic display of echo amplitude as a function of distance along one line of sight, or vector (Figure 2.1). A-mode was the first display mode to be used in ophthalmology (1). It is used in characterization of tissues such as intraocular tumors and vitreous hemorrhage. It is also widely used in biometric applications, such as axial length measurement and corneal pachymetry. B-mode, introduced in the late 1950s (2), refers to a display of two-dimensional cross-sectional images (Figure 2.2). These images provide representations of the anatomy of the eye and orbit that have proven useful in diagnosis of a broad spectrum of disease states. A- and B-mode systems may be found in instruments dedicated to one function only or may be combined in a single instrument. A-mode displays may be generated using a special purpose A-mode transducer or may be generated from individual vectors comprising a B-mode display.

Although the basic physical principles discussed in Chapter 1 underlie the operation of all ultrasonic systems, it is electronic and computer technology that translates these principles into practical clinical instruments. The electronic and computer components of modern ultrasound scanners are used to generate ultrasonic pulses, process echoes, and display images and information. This chapter discusses the electronic components used in each of these stages for generation of A- and B-mode images and how the characteristics of the individual components influence the quality of the resulting images. It also discusses the means of recognizing and eliminating misleading results stemming from improper system adjustment. Throughout this chapter, emphasis is given to the overall quality of an ultrasonogram in terms of three parameters: resolution, sensitivity, and dynamic range. Spatial resolution, defined in Chapter 1, refers to the ability to distinguish two nearby reflectors. Temporal resolution, also to be considered here, refers to the ability to visualize tissue changes occurring over time. Sensitivity refers to the weakest reflector that can be detected in a displayed ultrasonogram. Dynamic range describes the spread of echo amplitudes that can accurately be portrayed in an ultrasonogram.

In addition to A- and B-modes, this chapter will describe modalities less commonly used or more recently introduced in ophthalmology, including Doppler, M-mode, and swept-mode.

SYSTEM COMPONENTS

The ultrasound system, schematically represented in Figure 2.3, consists of the following components:

Transducer/probe

Servo (for B-mode systems)

Pulser

Receiver

Scan converter and display

The trend for system design in modern instrumentation is toward integrated digital components. By placing the previously mentioned functions on a single computer board, the system becomes less expensive, more reliable, and easier to repair. Because of this integration, not all of the components are represented as stand-alone devices, but we shall consider them as functionally separate entities.

PROBE

In B-mode systems, a mechanism is needed to sweep the ultrasound beam across a scan plane. In mechanical sector

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scan systems, which dominate in ophthalmology, the transducer is enclosed within a sealed, fluid-filled housing with an acoustically transparent cap at one end. In operation, the transducer is pivoted over an angle of 45 to 60 degrees at a rate of several times per second. This refresh rate, called the frame rate, is typically about 10 Hz, but in some instruments rates of 30 Hz or more have been attained.

Figure 2.1. A-scans consist of a plot of echo amplitude as a function of range. This figure shows a typical A-scan along the axis of a normal eye performed in contact mode through the eyelid. Peaks correspond to eyelid (L), cornea (C), anterior (AL), and posterior lens (PL) surfaces and retina (R).

SERVO

The servo is a device that controls the motion of the transducer within the probe and registers the orientation of the transducer at each moment of time. The servo controls a motor incorporated within the probe, and, as the transducer moves, the servo continually monitors its position. Each scan frame consists of a fixed number of vectors (typically 256) that are evenly spaced within each scan frame. As the motor sweeps the transducer, the servo monitors its position and issues signals to the pulser and other components such that pulse/echo vectors are acquired at appropriate positions.

Figure 2.2. Figure 2.1 illustrates the difficulty in establishing context for an A-scan seen in isolation. This

B-mode image shows the vector from which the A-scan in Figure 2.1 was derived but in the context of the ocular anatomy revealed in B-mode.