Ординатура / Офтальмология / Английские материалы / Handbook of Optical Coherence Tomography_Bouma, Tearney_2002

Figure 5 Different scanning protocols for OCT imaging. There are numerous scanning protocols that can be used for OCT imaging depending upon the imaging engine and application. Cross-sectional images may be acquired using either depthor transverse-priority scanning. In depth-priority scanning, axial scans are acquired at successive transverse positions. In transverse-priority scanning, the beam is scanned rapidly in the transverse direction and light from successive axial depths (or ranges) is detected. Transverse-priority scanning in two transverse dimensions may be used to acquire en face images at a given depth.

optical reflection or backscattering is mapped to different colors. In this image, the intensity of the optical signal is mapped to the color scale using the standard ‘‘rainbow’’ order of colors. The highest backreflection or backscattering is represented by red and white and typically corresponds to 100 dB of the incident signal, whereas the lowest backscattering is represented by blue and black and typically corresponds to 100dB of the incident signal. This image demonstrates that the use of false color can improve the differentiation of different structures. In contrast to gray-level

Figure 6 Gray-scale OCT image. Example of an OCT image of the fovea region of retina in a normal human subject. The image shows the differentiation of retinal layers that is possible using gray-scale display. The logarithm of the backscattered or backreflected signal is mapped into the gray scale. Typical ophthalmic images span approximately a 40–50 dB dynamic range. The ability of the eye to differentiate different gray levels is limited, and monitors support only 8-bit gray levels, so the fidelity of the image formation is lost.

monitors, color monitors can have 24-bit color levels, and the human eye can differentiate millions of different colors. The principal disadvantage of using false color display is that it can produce artifacts in the image. If the signal intensity is changed, this produces a color shift of structures in the image. Thus careful normalization of signal levels is required. In addition, different signal levels in the image are mapped to different colors that do not necessarily correspond to different physical structures.

The images of the retina presented in Figs. 6 and 7 (see color plate) are examples of imaging media that have weak reflections. Figure 8 shows a gray-scale OCT image of the human ectocervix in vitro as an example of an image in a highly scattering tissue. In a highly scattering material or tissue, light is rapidly attenuated with propagation depth, resulting in a gradation of signal in the image. Also, there can be significant levels of speckle noise or other noise arising from the microstructural

Figure 7 False color OCT image. Example of an OCT image of the fovea region of the retina in a normal human subject. The image shows the differentiation of retinal layers that is possible using false color display (compare to gray-scale display of Fig. 6). The logarithm of the backscattering or backreflected signal is mapped to a false color scale. The retinal pigment epithelium, choroid, and retinal nerve fiber layers are visible as highly backscattered red layers. Typical ophthalmic images span a 40–50 dB dynamic range. Because the eye can differentiate many more colors than gray levels and monitors can support 24-bit colormaps, false color improves the ability to differentiate structures. (See color plate.)

Figure 8 Gray-scale image in scattering tissue. Example of an OCT image of the ectocervix in vitro that illustrates features common to many scattering materials and tissues. These images show strong attenuation of the signal with depth as well as speckle noise. They are more prone to display artifacts than OCT images in weakly scattering tissues. Images in scattering materials or tissues are often displayed using a gray scale. Note the presence of glandular (g) structures that have different backscattering properties.

features of the material or specimen. These image properties can produce display artifacts if the image is displayed using a false color scale. Thus it is more common to use a gray scale to display OCT images in highly scattering materials or tissues.

It is important to note that although the tomographic image represents the true dimensions of the structure (correcting for index of refraction and refraction effects) being measured, the coloring of different structures represents different optical properties and not necessarily different tissue morphology. Thus, OCT images should be interpreted analogously to conventional histopathology. A comprehensive discussion of the theory of OCT imaging is presented in Chapter 2.

1.7PIXEL DENSITIES AND IMAGE ACQUISITION TIME

The pixel density of an OCT image is determined by the image acquisition conditions and analog-to-digital (A/D) sampling parameters. Assuming that each axial scan covers a depth Lz, then if the axial resolution is z the axial scan data should be sampled at a density of two times the resolution or higher. Thus it is desirable to have at least Nz ¼ 2Lz= z image pixels in the axial direction. The number of pixels in the transverse direction is determined by the number of axial scans, Nx, used to construct the OCT image. For optimum resolution, the number of transverse pixels should be chosen according to the transverse resolution x. Thus if an OCT image with a transverse dimension of Lx is required, the optimum resolution image should have Nx ¼ 2Lx= x pixels. However, in most cases this is a preclusively high number of transverse pixels, and OCT images are undersampled in the transverse direction. With improvements in technology, the pixel dimensions of OCT images have been steadily increasing. Early ophthalmic OCT imaging was performed with 100 transverse pixels and 500 longitudinal pixels or 50k image sizes [25], whereas now ultrahigh resolution OCT images can have more than 1000 transverse pixels and 1500 longitudinal pixels, resulting in megapixel image sizes [22].

Achieving rapid image acquisition time is important for imaging in vivo specimens and for clinical applications. Rapid imaging is necessary to minimize the image distortion produced by motion. In addition, the time available to perform the examination is limited, and it is desirable to maximize the amount of data obtained. The imaging time is directly related to the detection sensitivity, because performing imaging more rapidly (i.e., increasing the noise equivalent bandwidth of the detection) results in reduced signal-to-noise performance in the image. At the same time, in order to distinguish weak reflections from different intraocular structures in ophthalmic imaging or to achieve sufficient image penetration in scattering tissues that are strongly attenuating, it is necessary to image with a sufficient signal-to-noise ratio.

The signal-to-noise ratio may be improved by using higher incident optical power. However, for clinical applications the maximum permissible light exposure is determined by safety standards. For eye and skin exposure the American National Standards Institute (ANSI) has developed criteria for determining the maximum permissible exposure [26]. Unfortunately, no systematic or accepted criteria exist that govern the safe exposure of epithelial surfaces such as those in the gastrointestinal, pulmonary, urinary, or reproductive tracts that are exposed in internal body imaging applications. Thus, the establishment of exposure standards for scanned focused spot illumination of the type performed in OCT imaging (as well as in confocal microscopy) remains an important scientific and regulatory problem.

The image acquisition time is determined by a combination of incident power and signal-to-noise ratio that is required to achieve images of sufficiently high quality for given applications. Image acquisition time increases in proportion to the size of the image and the number of transverse resolution elements or pixels. The image acquisition time is given by the amount of time it takes to perform each axial scan times the number of axial scans in the image, T ¼ Nxfs, where Nx is the number of transverse pixels in each tomograph and fs is the repetition rate of the axial scanning. This also determines the velocity of the axial scanning, because the distance Lz in the axial direction must be scanned in a time of 1=fs or less. The scan velocity is thus

vs ¼ Lz=fs.

If higher resolution imaging is desired in the transverse direction, i.e., a greater number of transverse pixels are desired, then the image acquisition time increases proportionally. Conversely, if very low resolution imaging is performed—if, for example, only topographic information is required—then the number of transverse pixels may be reduced, resulting in proportionally faster image acquisition. In the longitudinal direction, the image acquisition time scales directly as the depth or length of the longitudinal axial scan necessary to image the desired structure.

1.8IMAGE CONTRAST PENETRATION DEPTHS

The detection sensitivity determines the imaging performance of optical coherence tomography for different applications. It is helpful to consider OCT imaging in two limiting cases: (1) imaging in media with very weak scattering and (2) imaging in highly scattering media. When OCT is performed in very weakly scattering media, the imaging depth is not strongly limited, because there is very little attenuation of the incident beam. Instead, the sensitivity of the OCT detection establishes a limit on the smallest signals that can be detected. One example of this is in optical data

storage applications that encode information using reflections from index mismatch [27]. In these applications it is desirable to use as small an index mismatch as possible. If an optical reflection is generated as the result of an index mismatch n, then the magnitude of the intensity reflection is n2. The high sensitivity of OCT means that extremely small reflections corresponding to small index mismatches can be measured. For example, if the detection sensitivity is 100 dB, index mismatches in the range of n ¼ 10 4–10 5 can be detected. The application of OCT for data storage is described in more detail in Chapter 14.

Another example of OCT imaging in structures with weak backscattering is ophthalmic imaging. Figures 6 and 7 show examples. In ophthalmic imaging, high sensitivity is essential in order to image structures such as the retina that are nominally transparent and have very low backscattering. In ophthalmic applications, the image contrast in OCT images arises because of differences in the backscattering properties of different tissues. Structures such as the retinal pigment epithelium (RPE) or retinal nerve fiber layer can be differentiated from other structures because of their different scattering amplitudes. Typical retinal images have signal levels within 50 to 100 dB of the incident signal. For retinal imaging, the ANSI standards govern the maximum permissible light exposure and set limits for both the sensitivity of OCT imaging and OCT imaging speeds [25,26,28]. The application of OCT for ophthalmic imaging is described in more detail in Chapters 17 and 18.

The second limiting case of OCT imaging is imaging in highly scattering media. Figure 8 shows an example. In this case, the detection sensitivity determines the maximum depth to which imaging can be performed. Most biological tissues are highly scattering. OCT imaging in tissues other than the eye became possible with the recognition that the use of longer optical wavelengths, for example, 1:3 m compared to 800 nm, can reduce scattering and increase image penetration depths [29– 32]. Figure 9 shows an example of OCT imaging in a human epiglottis in vitro comparing 800 nm and 1300 nm imaging wavelengths. The dominant absorbers in most tissues are melanin and hemoglobin, which have absorption in the visible and near-infrared wavelength range. Water absorption becomes appreciable for wavelengths approaching 1.9–2 m. In most tissues, scattering at near-infrared wavelengths is one to two orders of magnitude higher than absorption. Scattering decreases for longer wavelengths, so OCT image penetration increases [33]. For example, if a tissue has a scattering coefficient in the range of 40 cm 1 at 1300 nm, then the round-trip attenuation from scattering alone from a depth of 3 mm is e 24 or 4 1011. Thus, if the detection sensitivity limit is 100 dB, backscattered or backreflected signals are attenuated to the sensitivity limit when the image depth is 2–3 mm. Because the attenuation is exponential with depth, increasing the detection sensitivity by one order of magnitude would not appreciably increase the imaging depth.

The mechanisms of OCT image contrast in scattering media have been investigated and related to the optical properties of the medium [34]. The image contrast mechanisms for OCT are somewhat analogous to those of ultrasound. In order to visualize an internal feature in a highly scattering medium, it is necessary to detect light that is backscattered or backreflected from the internal feature. First, the OCT imaging beam is incident on the scattering medium and is attenuated by scattering and absorption during its forward propagation into the medium. The light is backscattered or backreflected by the internal feature and is further attenuated by scat-

Figure 9 Wavelength dependence of OCT image penetration depth. Example of OCT images of human epiglottis in vitro acquired with light sources at 850 and 1300 nm. Attenuation from scattering is reduced by using longer wavelengths. Superficial glands (g) can be visualized at both wavelengths, but the underlying cartilage (c) can be imaged only with 1300 nm light. With 100 dB detection sensitivity, image penetration of 2–3 mm is possible in most scattering tissues. The ability to perform OCT imaging in scattering tissues opened up a wide range of medical applications. (From Ref. 32.)

tering and absorption during its reverse propagation out of the medium. Thus, the contrast for OCT imaging in scattering media is determined by a combination of attenuation from scattering and absorption during propagation and backscattering from the internal feature that is being imaged. The situation is different for OCT imaging in weakly scattering media, where there is negligible attenuation from scattering and absorption during propagation. In this case, image contrast is determined by the backscattering properties of the internal features that are being imaged. A detailed discussion of tissue optical properties and mechanisms of OCT image contrast is presented in Chapter 16.

1.9RESOLUTION AND IMAGE PENETRATION DEPTHS OF OCT AND ULTRASOUND

Two of the most important parameters for characterizing imaging performance are image resolution and imaging depth. There are interesting comparisons between OCT, ultrasound, and microscopy as shown in Fig. 10. The resolution of ultrasound imaging depends directly on the frequency or wavelength of the sound waves that are used [1–5]. For typical clinical ultrasound systems, sound wave

Figure 10 Resolution and penetration of ultrasound, OCT, and confocal microscopy. In ultrasound imaging, both image resolution and penetration are determined by the sound wave frequency. High axial resolutions are possible, but transverse resolutions are usually coarser. Ultrasonic attenuation limits penetration to a few millimeters. The image resolution in OCT is determined by the coherent length of the light source, and resolutions ranging from 1 to 15 m can be achieved. In most scattering tissues, image penetration is limited to 2–3 mm. Confocal microscopy can have submicrometer transverse resolution, but axial resolution is coarser. The image penetration of confocal microscopy is typically limited to a few hundred micrometers in scattering tissues. Together these imaging modalities provide complementary performance.

frequencies are in the 10 MHz regime and yield spatial resolutions of up to 150 m. Ultrasound imaging has the advantage that sound waves at this frequency are readily transmitted into most biological tissues and therefore it is possible to obtain images of structures up to several tens of centimeters deep within the body. High frequency ultrasound has been developed and investigated extensively in laboratory applications as well as in some limited clinical applications. Axial resolutions of 15–20 m and higher have been achieved with frequencies of 100 MHz. It should be noted that the transverse resolution of ultrasound is governed by the dimension of the focused ultrasound beam. Although the sound frequency governs the focused spot size, it is difficult to achieve spot sizes on the order of the wavelength, and thus transverse resolutions are not as fine as axial resolutions. Another important limitation of high frequency ultrasound is that high frequencies are strongly attenuated in biological tissues, with attenuation increasing approximately in proportion to the frequency. Thus, high frequency ultrasound imaging is limited to depths of only a few millimeters.

Figure 12 The systems perspective of OCT. OCT is a modular technology consisting of both hardware and software. The major submodules in the OCT system include the imaging engine, light source, delivery system and optical probes, computer control, and image processing. OCT draws upon a wide range of component technologies.

[35–38]. A comprehensive discussion of Doppler imaging using OCT is presented in Chapter 8. Polarization-sensitive detection techniques have been demonstrated using a dual channel interferometer [39–41]. These techniques permit imaging of the birefringence properties of structures. Collagen and other tissues are strongly birefringent, and polarization-sensitive OCT can be a sensitive indicator of the tissue microstructural organization. Conversely, polarization diversity or polarizationinsensitive interferometers can also be built using well-established polarization diversity detection methods from coherence heterodyne optical communication. A detailed discussion of techniques for OCT imaging and measurement of polarization is presented in Chapter 9. Finally, other imaging engine approaches have been demonstrated that are based on spectral analysis of broadband light sources as well as tunable narrow linewidth sources [42]. These imaging engines have the advantage that they do not require scanning an optical delay, and thus no moving parts are required. Discussions of these types of imaging techniques are presented in Chapters 12 and 13.

The short coherence length light source determines the axial resolution of the OCT system. For research applications, short-pulse lasers are powerful light sources for OCT imaging because they have extremely short coherence lengths and high output powers, enabling high resolution, high speed imaging. Many studies have been performed using a short-pulse Cr4þ : forsterite laser. This laser produces output powers greater than 100 mW with ultrashort pulses of less than 100 fs at wavelengths near 1300 nm. By using nonlinear self-phase modulation in optical fibers, the spectral output can be broadened to produce bandwidths sufficient to achieve axial resolutions of 5–10 m [43]. Image acquisition speeds of several frames per second are achieved with signal-to-noise ratios of 100 dB using incident powers in the 5–10 mW

range. Short-pulse Ti:Al2O3 lasers operating near 800 nm have also been used to achieve high resolutions. Early studies demonstrated axial resolutions of 4 m [44]. Recently, using Ti:Al2O3 laser technology, which generates pulse durations of 5 fs and bandwidths of over 300 nm, axial resolutions of 1 m were achieved [22]. For clinical applications, compact superluminescent diodes or semiconductor-based light sources have been used. Although these sources have not yet reached the performance levels of research systems, they are compact and robust enough to be used in the clinical environment. Ophthalmic OCT systems have employed commercially available (EG&G Optoelectronics) compact GaAs superluminescent diodes that operate near 800 nm and achieve axial resolutions of 10 m with output powers of a few milliwatts. Commercially available (AFC Technologies Inc.) sources based on semiconductor amplifiers operating at 1:3 m can achieve axial resolutions of 15 m with output powers of 15–20 mW, sufficient for real-time OCT imaging. With further development, other types of low coherence light sources based on rare-earth- doped fibers, Raman conversion, and other nonlinear conversion processes should also become feasible for clinical OCT systems. A detailed discussion of low coherence light sources is presented in Chapter 3.

In the reference delay scanning embodiments of OCT systems, the optical delay scanner determines the image acquisition rate of the system. The earliest scanning devices were constructed using galvanometers and retroreflectors [25,45]. These systems have scan ranges of up to a centimeter or more and have the advantage of simplicity. However, scan repetition rates are limited to approximately 100 Hz. A novel technique using PZTs to stretch optical fibers has been developed that can achieve scan ranges of a few millimeters with repetition rates of 500 Hz or more [46,47]. Many studies were performed using a high speed scanning optical delay line based on a diffraction grating phase control device [48]. This device is similar to pulse-shaping devices that are used in femtosecond optics [49]. The grating phase control scanner is attractive because it achieves extremely high scan speeds and also permits the phase and group velocity of the scanning to be independently controlled. Images of 250–500 transverse pixels can be produced at several frames per second [50]. A detailed discussion of high speed scanning techniques is presented in Chapter 4.

The OCT beam delivery and optical probes can be designed for specific applications. Because OCT imaging technology is based on fiber optics, it can be easily integrated with many standard optical diagnostic instruments, including instruments for internal body imaging. OCT imaging has been integrated with a slit lamp and fundus camera for ophthalmic imaging of the retina [25,51]. The OCT beam is scanned using a pair of galvanometric beam-steering mirrors and relay-imaged onto the retina. The ophthalmic OCT imaging instrument is relatively complex because it must image the retina (fundus) en face while showing the position of the tomographic scanning beam. The ability to register OCT images with en face features and pathology is an important consideration for many applications. Similar principles apply for the design of OCT using low numerical aperture microscopes. Low numerical aperture microscopes have been used for imaging in vivo developmental biological specimens as well as for surgical imaging applications [52–55].

Closely related beam delivery systems include forward imaging devices, which permit the delivery of a oneor two-dimensional scanned beam. Rigid laparoscopes are based on relay imaging using Hopkins-type relay lenses or graded index rod