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

Figure 2.3. Schematic representing electronic components comprising a B-mode system. These include the probe (containing the transducer and mechanical scan mechanism), the pulser/receiver (which excites the transducer and amplifies echoes), the servo (which controls and monitors transducer orientation), the scan converter (which digitally formats echo data for display), and the display device.

PULSER

The ultrasonic pulser repeatedly “shock excites” the transducer with short voltage pulses applied across the electrodes of its piezoelectric element. Each excitation results in the generation of an ultrasonic pulse. The pulse repetition frequency (PRF) must be low enough to allow all returning echoes to be received by the transducer before the next pulse is generated. If we consider the distance to the optic chiasm (the longest distance that we would need to consider) to be approximately 6 cm, then we can determine the two-way travel time for an acoustic pulse to be 2 × 0.06 m ÷ 1,540 m/sec = 0.078 msec. Typically, a 1 KHz pulse repetition frequency is used so that pulses are generated at 1-msec intervals, much longer than the time for pulse and echo travel from the deepest portion of the orbit.

The nature of the ultrasonic pulse is an important factor in determining the characteristics of the emitted acoustic pulse, and hence, of the attainable resolution of the ultrasound system. The characteristics of the pulse that are significant include its form, duration, and amplitude. The most common pulse form used in A- or B-mode imaging is a negative spike, although other modes, such as monocycle (a single sine-wave) are sometimes used. Long excitation pulses result in correspondingly long ultrasonic pulses and concomitantly poor axial resolution. Thus, negative spike impulses are normally designed to be of short duration. Unlike negative spike impulses, monocycles can be tuned to a specific frequency that may or may not correspond to the natural harmonic frequency of the transducer. Monocycles can thus be used to alter

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the output of a given transducer to frequencies that are somewhat higher or lower than its natural frequency. The amplitude of the excitation pulse is important in achieving adequate sensitivity. Typically, an amplitude of 100 to 400 volts will result in ultrasonic pulses large enough to produce detectable echoes from weakly reflecting interfaces, such as the vitreoretinal interface. Although higher voltages provide higher amplitude pulses, and, consequently, increased sensitivity, extremely high voltages can result in distortion of the acoustic pulse (e.g., ringing) that results in decreased axial resolution. In some systems, access to pulser characteristics, such as damping and energy, allows the excitation pulse to be varied between extremes of a short, low-amplitude pulse (providing high resolution and low sensitivity) and a long, high-amplitude pulse (providing low resolution and high sensitivity). In instruments where pulse characteristics are accessible to the user, the clinician can determine the combination of resolution and sensitivity most advantageous in a given situation.

RECEIVER

Following transducer excitation, echoes from intraocular and orbital structures impinge upon the transducer, which generates small voltages that are proportional to echo amplitude. These voltages must be processed by the amplifier and associated electronics before they can be displayed in a useful format. The required operations are carried out by the electronic receiver, whose most important function is amplification of the minute voltages generated by the transducer. Gains of 100 (40 dB) or more are needed to raise the amplitudes of these signals from their initially low levels (e.g., 1 millivolt) to levels that are compatible with display requirements. The receiver's functions include a limiter to prevent damage from the high voltage excitation spike produced by the pulser, sometimes a low-noise pre-amplifier (which, in some systems, is incorporated in the probe) to boost gain, and a variety of other functions, possibly including time-gain control (TGC), compression, noise reject, and envelope detection.

A factor of prime importance when regarding receiver gain is saturation. This occurs whenever the amplified pulse reaches the maximum level (e.g., 5 volts) that the amplifier can supply. If an input signal is too large (or if the gain is too high), saturation occurs and can eliminate clinically significant information regarding echo strength (Figure 2.4). Saturation is always encountered at the beginning of an A- or B-mode display, when a part of the large excitation pulse is picked up by the amplifier. This so-called “main bang” causes a dead space directly in front of the transducer and obliterates echoes from close objects. For this reason, the region of the main bang is not displayed in most scan systems and is the reason that all transducers require some sort of a “stand off” between the transducer and the tissue of interest.

Figure 2.4. All amplifiers have a saturation level above which signals are truncated or otherwise distorted. Reject levels, which may be user-controllable, allow low-level signals (such as noise) to be suppressed. The figure demonstrates the effects of saturation and reject on highand low-level signals.

Electronic noise limits the detection of small echoes and can present severe problems with large bandwidth amplifiers. Noise consists of small, random voltage variations that arise because of statistical fluctuations of electrons in system components. At high gains the resultant noise signals appear as “grass” in an ultrasonogram and obscures low amplitude echoes. Some ultrasound units provide reject controls that establish a threshold level that can be set to prevent noise from being displayed. Reject controls, if improperly adjusted, can suppress the display of small or modest echoes that may be diagnostically significant (Figure 2.4).

Amplifiers often include filters that pass signals within only a certain frequency range (usually surrounding the transducer center frequency). This frequency bandpass operation reduces noise from extraneous sources.

Saturation and noise combine to determine the dynamic range of an amplifier. Dynamic range is defined as the spread of input signal amplitudes that result in meaningful output signals. The largest useful input signal is one that just causes saturation; the smallest useful input signal is one that yields an output just above the amplifier noise level (or reject level). Typical amplifiers have dynamic ranges of 60 dB, that is, the maximum input signal is one million times larger than the minimum input signal. This range is sufficient for accurate displays of tissue echoes.

TGC is a form of amplification in which gain increases as a function of range. This feature is designed to compensate for the reduction in echo amplitude that occurs with depth as a result of attenuation of acoustic energy with depth. As noted in Chapter 1, absorption causes a progressive loss in the strength of a propagating ultrasonic pulse. Absorptive attenuation of echo strength increases exponentially with distance. Compensation for this effect involves increasing amplifier gain with time, t, in a manner which is the inverse of absorptive

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decay. The appropriate gain is thus of the form of e?ct, where ? is the absorption coefficient and c is the velocity of ultrasonic propagation. Although TGC is useful, it cannot compensate exactly for absorption in the eye and orbit because absorption coefficients differ greatly in specific ocular and orbital tissues. Most instruments, therefore, do not use the function shown previously, but rather let the user adjust gain arbitrarily over a series of “gates,” that is, successive nonoverlapping range windows. An example of a TGC-modified image is provided in Figure 2.5.

Envelope detection, or demodulation, is used to convert the positive and negative voltages (and nulls) that are present in the raw radiofrequency (RF) echo data to positive signals, sometimes referred to as the video signal or envelope. This operation, shown schematically in Figure 2.6, includes a rectification stage, in which all negative voltages are converted to positive values, followed by a suitable low-pass filter operation, which is designed to smooth out the nulls in the rectified RF waveform, without excessive blurring. Demodulation is a crucial step for formation of images, because A- and B-mode images can represent positive values only. It is the signal envelope that is generally used for formation of A-mode and B-mode images.

Figure 2.5. Time gain control (TGC) consists of manipulation of gain as a function of range, thus allowing enhancement of signals that would otherwise be lost as a result of attenuation. Top: B-scan of an eye and orbit without use of TGC, as indicated by the flat TGC curve below the image. Bottom: Same image with TGC

applied, demonstrating enhancement of the orbit.

Figure 2.6. This schematic illustrates stages in processing of echo data. Radiofrequency (RF) data represent the positive and negative voltages associated with the positive and negative pressures of the acoustic waves but are inconvenient for display purposes. The first stage of processing is rectification, in which the negative voltages are rendered positive. Center: The rectified data, however, are choppy due to the nulls in the RF waveform. The rectified data are therefore smoothed using a low-pass filter, as shown in the lowermost waveform. This waveform is termed the “envelope” and is used for A- and B-mode displays.

Compression is an operation that reduces the contrast between high and low amplitude echoes. This operation is important because display devices have a dynamic range that is far less than the amplitude range of linearly amplified echoes. Logarithmic amplification is one such compression scheme. In a standard (linear) amplifier, the gain, G, is independent of the input signal level, and an output voltage is equal to G times the corresponding input pulse. In a logarithmic amplifier, the

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output voltage is proportional to the logarithm of the input voltage. As shown in Figure 2.7, this type of amplification yields a large gain for small signals and a small gain for large signals. Thus, a logarithmic characteristic reduces differences in echo levels and permits a wider range of input signals to fall between the noise and saturation levels of an amplifier or digitizer. Logarithmic amplification can be used effectively in B-mode systems, where it compresses the large range of input echo signals to the smaller range of brightness levels that can be presented on digital displays. Other nonlinear functions have also been used to advantage in ultrasonography. An S-shaped function, for example, provides the greatest dynamic range at intermediate signal levels and has proven useful in A-mode examinations, as used in “standardized” A-scan ultrasonography (3). However, such nonlinear amplification makes it difficult to recognize relative echo amplitudes and to compare clinical echo patterns with those obtained in standard systems. Characteristic A-mode patterns can be easily distorted, unless great care is exercised in adjusting all system parameters.

Figure 2.7. Logarithmic amplification may be used to extend dynamic range by reducing contrast between bright

and faint echoes. This allows display of faint echoes without causing saturation of bright echoes.

Demodulation converts RF data into a convenient format for display but does this at the sacrifice of some of the information content inherent in the original echo data. If we recall that the reflection coefficient is defined as (Z2-Z1)/(Z1 + Z2), where (as in Chapter 1) Z refers to acoustic impedance, then we can see that the

interface reflection, though of the same magnitude, will change in sign when going from a medium of high to low impedance (Z1 > Z2), versus low to high

impedance (Z1 < Z2). This effect is illustrated in Figure 2.8. Thus, RF data can provide information about the direction and magnitude of impedance change across

an interface, whereas the demodulated signal provides information regarding the magnitude only. RF data processing is now the preferred approach for extracting

subtle characteristics regarding tissue microstructure, and many manufacturers of sophisticated ultrasound systems used in specialties outside of ophthalmology

now incorporate RF signal capture.

Figure 2.8. Radiofrequency data derived from indicated (horizontal line) vector segment of 50-MHz scan of cornea. Note that the phase of the posterior corneal surface echo (right arrow) is inverted in relation to that of the anterior surface echo (left arrow). This effect arises from the opposite signs of the reflection coefficients of the surfaces, which is positive anteriorly [(1640-1540)/(1640 + 1540) = 0.031] and negative posteriorly

[(1540-1640)/(1540 + 1640) = -0.031], where the speed of sound in cornea is 1,640 msec-1, and the saline

coupling medium and aqueous are 1,540 msec-1.

SCAN CONVERTER AND DISPLAY

Early ultrasound instruments used analog devices, such as oscilloscopes or television monitors, as display devices. This has been virtually superseded by digital displays. An analog-to-digital converter (ADC) is needed to use digital display devices.

Analog-to-digital conversion involves transformation of a continuous signal into a discrete binary representation of the kind used in computers. The ADC samples the output of the amplifier at a specific rate, which is normally at least twice the highest frequency component present in the signal. Although the original analog signal is unconstrained in its amplitude values (within the dynamic range of the amplifier), ADC limits the range and number of levels of data that can be acquired.

An 8-bit ADC, for instance, can store up to 28 = 256 voltage levels. If a voltage level exceeds the range of the ADC, then it is “clipped.” At present, 8-bit digitizers are most common in ophthalmic ultrasound systems, but deeper bit-depth components are available. A 12-bit digitizer, for instance, can represent 4,096 levels. As deep bit-depth digitizers become incorporated in instruments, the need for compression, which always involves some information loss, will decrease.

For digital display of ultrasound data, a special type of ADC, the scan converter, is used. The scan converter

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holds a series of memory locations that map directly to pixels of the digital display. The primary function of the scan converter is transformation of B-mode data into an image representation in digital memory. This operation includes a remapping of every displayed pixel to specific scan vectors and ranges, such that proper image geometry is obtained. With each successive scan frame, scan converter memory is overwritten and the new information displayed on the monitor. When the freeze frame button of the scanner is pushed, the last update of the scan converter memory is continually shown, allowing prolonged examination of individual frames and storage as computer files. Some systems also include a cine-loop feature in which not one but several successive frames, or even short segments of real-time display, are stored in digital memory. This feature allows segments, usually a few seconds in length, of the real-time exam to be reviewed and individual frames to be chosen and stored.

Most systems also include a digital-to-analog converter, whose purpose is to convert the digital image information into standard video formats, such as NTSC or

Figure 2.10. Typical A-scan printout for axial length measurement. Left to Right: Echoes represent cornea, anterior and posterior lens, and retina. Note automatic determinations of anterior chamber and lens thickness, in addition to axial length.

The most common use of A-scan is for axial length measurement (Figure 2.10). These units generally operate at a 10-MHz center frequency with weakly focused transducers: a 5-mm aperture and 20-mm focal length are typical. A-scan probes may be either handheld or attached to a slit-lamp type mount. During operation, excitation pulses are emitted at a high PRF, such that many waveforms are captured per second. A-scan units typically incorporate pattern recognition software that detects the shape of A-scans and automatically acquires those that meet criteria expected when the probe is properly aligned. These data can then be analyzed for measurement of axial length, anterior chamber depth, and lens thickness. Numerous formulas (4, 5, 6, 7) have been developed for computation of lens implant power, based on these measurements and keratometry readings. These computations must take into account speed of sound in special instances, such as a cataractous lens, aphakia, and so forth. A-scan instruments also allow measurements to be made manually from the A-scan trace, which may be necessary in circumstances of unusual ocular anatomy.

Corneal pachymetry (derived from the Greek pachy, meaning “thick,” and metron, meaning “measure”) (8) is a form of A-scan, whose aim is to measure corneal thickness, which is, in principle, no different from axial length measurement. In the case of corneal pachymeters, a transducer frequency of 20 MHz and aperture less than 2 mm are typical, although units operating as high in frequency as 50 MHz are available. Data over a range of about 1.5 mm are plotted. Pachymetry probes, like axial length probes, can be either handheld or mounted. Pachymetry can also be performed intraoperatively in LASIK to measure the thickness of the residual stroma (9). A system whose bandwidth is sufficient for processing 20-MHz data can readily process 10-MHz data as well. In fact, some A-scan instruments incorporate both types of probes. In such a case, the manufacturer will provide excitation pulses and filters suitable for each transducer configuration, along with software suitable for performing the appropriate biometric determinations.

A special purpose A-scan probe is used in an ophthalmic ultrasound technique called standardized echography (10). Standardized echography units incorporate an unfocused 8-MHz A-scan probe and a focused 10-MHz B-scan probe. In practice, the A-scan probe is placed in direct contact with the globe and used for identification of intraocular tumors and differentiation of retinal detachment from vitreous membrane, among other applications. Standardized echography units use an amplification curve that is S-shaped (a form of compression) to place the dynamic range of the system in the mid-range of echo amplitudes. Calibration is performed by measurement of the reflections produced by a standard tissue model, thus linking the level of the S-shaped amplification curve to known decibel levels.

A-scans can also be derived from B-mode images. If we recall that a B-mode image is simply a composite of many consecutive spatially offset vectors, then it is apparent that we can readily obtain a plot of the amplitude envelope as a function of range along any of the vectors comprising the B-mode image. For this to be quantitative, however, it is necessary to compensate for the effects of gain, TGC, and any other transformations used in generating pixels from echo data. An A-scan generated in this manner would have the advantage of providing quantitative information in the context of the two-dimensional B-scan. For axial length measurements, use of A-scans generated from B-mode data is generally not ideal, because scans are obtained (in contact mode) through the closed lid. Even if an immersion technique is used, the B-mode probe does not provide a fixation light to ensure that the scan is aligned with the visual axis. However, in special situations, for example, in eyes with irregular contours (staphyloma), A-scans derived from the B-mode image are the most useful, because they allow selection of the anatomically appropriate vector (11).

B-MODE SYSTEMS

B-mode systems (Figure 2.11) combine transducer scanning and signal processing to produce cross-sectional images of the eye and orbit. The quality of these images depends upon the factors already discussed for A-mode systems. However, there are additional electronic,

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mechanical, and acoustic considerations that should be understood for optimal clinical use and proper diagnostic interpretation. Furthermore, B-mode images are

susceptible to several types of artifacts that are readily recognized and that can often be eliminated. In addition to treating these topics, the following sections discuss useful image enhancement techniques and describe color-coding of pixels and three-dimensional imaging.

Figure 2.11. B-scan instruments incorporate a monitor for display of real-time gray-scale B-mode images. Some

B-mode instruments incorporate A-scan probes and scan modes suitable for axial length measurement as well. (Photo courtesy of Innovative Imaging, Inc.)

B-MODE IMAGE GENERATION

Conceptually, the simplest B-mode system uses a linear transducer scan motion, as shown in Figure 2.12. As the transducer is moved perpendicular to its beam axis, it emits acoustic pulses and receives echoes at a series of equally spaced positions. Each of these positions is referred to as a vector. Ultrasound systems are generally designed so that the distance between adjacent vectors is smaller than the transducer beam width in the focal plane (focal ratio × wavelength). If a larger distance is used, then the vectors are too sparse to guarantee that all anatomic structures will be captured in a single scan or sweep.

Figure 2.12. B-mode images are formed by scanning the eye by physical movement of the transducer. In the diagram, a simple linear transducer motion is used. The image is formed by slaving the display to the transducer position and range to each echo.

Although the scan motion is continuous, it is convenient to consider what happens at each vector location. Echo signals are processed, as in A-mode operation. However, the resulting video signals are not used to generate a plot of echo amplitude versus range, as in A-scan. Instead, these signals are used to regulate the intensity of the display. As described previously, this is accomplished by the scan converter, which is a rectangular array of digital memory locations having a one-toone correspondence with each pixel on the display device. In the case of a linear scan motion, there is a very simple relationship between these locations and the signal envelope. If we treat the scan converter memory as a rectangular array, then one axis corresponds to range, and the other to transducer position, with appropriate scaling factors in both cases. Because the scan converter may have more lines than vectors, interpolation is performed to fill in these otherwise empty memory locations. Thus, the display presents tissue reflectivity in terms of brightness as a function of time or, equivalently, distance from the transducer.

Thus, pixels indicate the two-dimensional position of reflective surfaces in the scanned tissue. In these images, sharply demarcated boundaries (e.g., the anterior lens surface) that generate well-defined A-mode echoes are displayed as distinct surfaces; acoustically homogeneous regions (e.g., the normal vitreous) are displayed as dark areas; acoustically heterogeneous areas (e.g., the orbital fat) that generate many closely spaced A-mode echoes are displayed as correspondingly speckled brightness patterns.

An inherent assumption in B-mode image generation is that of a constant speed of sound value throughout the entire image. This assumption facilitates generation of the image by allowing use of a single scaling factor to produce a geometrically correct image. However, the speed of sound varies throughout the eye, being significantly higher, for instance, in the cornea, sclera, and certain tumors than in the vitreous. This difference results in anamorphic distortion, which, although subtle, can introduce error into biometric determinations. Anamorphically corrected biometry can be performed by taking into account the depth and velocity of each tissue traversed, as well as vector location.

SCAN PATTERNS

B-mode systems can use various scan patterns, as shown in Figure 2.13. The most useful patterns are those in which the ultrasonic beam is perpendicularly aligned with reflective tissue surfaces. With perpendicular alignment, echoes travel directly back to the transducer rather

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than being redirected along a path that bypasses the transducer. Linear scan patterns can achieve perpendicularity only over small segments of curved ocular surfaces, such as those of the retina; therefore, they provide images of limited portions of the eye. Sector scan patterns are more compatible with these curved surfaces and allow echoes from large segments of posterior surfaces to be captured for B-mode presentations. Arc scan patterns permit perpendicular alignment over the anterior surfaces of the eye and lens. Of these patterns, the sector scan is the most widely implemented in ophthalmology. The sector scan is readily implemented in a sealed, compact probe, with a small footprint suitable for contact examination of the eye. Additionally, the sector scan involves the least physical motion of the transducer, and thus most easily achieves high scan repetition rates. The arc scan, however, is most suited for visualization of the anterior segment with very high frequency ultrasound (12,13). With a transducer frequency of about 40 MHz, examinations must be performed using an immersion technique with a lid speculum because of the effect of attenuation by the eyelid. In a noncontact exam, the arc scan is readily implemented and is ideal for visualization of the cornea and other anterior segment structures, as illustrated in Figure 2.14.