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

between OCT and ultrasound, it is helpful to consider the processes that govern OCT versus ultrasound imaging.

Ultrasound is an established clinical modality for internal body imaging [1– 5]. Ultrasound is used in many clinical applications, including imaging of internal organs, transluminal endoscopic imaging, and catheter-based intravascular imaging. In ultrasound, a high frequency sound wave is launched into the material or tissue being imaged by using an ultrasonic probe transducer. The sound wave travels into the tissue and is reflected or backscattered from internal structures that have different acoustic properties. Depending upon the frequency, significant attenuation of the sound wave may occur with propagation. The time behavior or echo structure of the reflected sound waves is detected by the ultrasonic probe, and the ranges and dimensions of the internal structures are determined from the echo delay.

In optical coherence tomography, measurements of distance and microstructure are performed using light that is backreflected and backscattered from microstructural features within the material or tissue [6]. For the purposes of illustration, it is possible to visualize the operation of OCT by thinking of the light beam as being composed of short optical pulses. However, it is important to note that although OCT may be performed using short-pulse light, most OCT systems operate using continuous wave short coherent length light. In addition, other OCT measurement approaches have been demonstrated that measure the spectral properties of low coherence light or use rapidly tunable narrow linewidth light.

The dimensions of structures can be determined by measuring the ‘‘echo’’ time it takes for sound or light to be backreflected or backscattered from the different structures at varying axial (longitudinal) distances. In ultrasound, the axial measurement of distance or range is referred to as A-mode scanning. The principal difference between ultrasound and optical imaging is that the velocity of light is approximately a million times faster than the velocity of sound. Because distances within the material or tissue are determined by measuring the ‘‘echo’’ time delay of backreflected or backscattered light waves, this implies that distance measurement using light requires ultrafast time resolution. Figure 1 compares the characteristic distance and time scales for light and sound propagation. The velocity of sound in water is approximately 1500 m/s, whereas the velocity of light is approximately 3 108 m=s. Distance or spatial information may be determined from the time delay of reflected echoes according to the formula T ¼ z=v, where T is the echo delay, z is the distance the echo travels, and v is the velocity of sound or light.

The measurement of distances or dimensions with a resolution on the 100m scale, which would be typical for ultrasound, corresponds to a time resolution of approximately 100 ns. One advantage of ultrasound is that echo time delays are on the nanosecond time scale and are well within the limits of electronic detection. Ultrasound technology has dramatically advanced in recent years with the availability of high performance and low cost digital signal processing technology. It is also important to note that because ultrasound imaging depends on sound waves, it requires direct contact with the material or tissue being imaged or immersion in a liquid to transmit the sound waves. In contrast, optical imaging techniques such as OCT can be performed without physical contact or

Figure 1 Echo time delay of light and sound. This figure shows the characteristic time and distance scales that govern ultrasound and optical ranging techniques. Optical coherence tomography (OCT) is analogous to ultrasound except that it performs imaging by measuring the echo delay of light rather than that of sound. The velocity of sound in water is 1500 ms=, whereas the velocity of light is 3 108 m=s. Because of this large difference in velocities, OCT and ultrasound use very different detection technologies.

the need for a special transducing medium. In applications such as ophthalmology, the use of noncontact imaging is important for reducing patient discomfort during examination. In endoscopy or bronchoscopy, imaging without a soundtransducing medium means that contact with or occlusion of the lumen is not required.

The echo time delays associated with light are extremely short. For example, the measurement of a structure with a resolution on the 10 m scale, which is typical in OCT, corresponds to a time resolution of approximately 30 fs. Direct electronic detection is not possible on this time scale. Thus, OCT measurements of echo time delay are based on correlation techniques that compare the backscattered or backreflected light signal to reference light traveling a known path length. Although OCT imaging is analogous to ultrasound, the core technology upon which it is based is quite different.

1.3MEASURING ULTRAFAST OPTICAL ECHOES

The concept of using high speed optical gating to perform imaging in scattering systems such as biological tissues was first proposed by M. Duguay 30 years ago [7–9]. Duguay demonstrated an ultrafast optical Kerr shutter to photograph light in flight. The Kerr shutter can achieve picosecond or femtosecond resolution and operates by using an intense ultrashort light pulse to induce birefringence (Kerr effect) in an optical medium placed between crossed polarizers. Duguay recognized that optical scattering limits imaging in biological tissues and that a high speed shutter could be used to gate out unwanted scattered light and detect light echoes from internal structures. This technology could see through tissues and noninvasively image internal biological structures.

The principal disadvantage of the Kerr shutter is that it requires high intensity laser pulses to induce the Kerr effect and operate the shutter. An alternative approach for high speed gating is to use second harmonic generation or parametric conversion. The objective or specimen being imaged is illuminated with short pulses, and the backscattered or backreflected light is upconverted or parametrically converted with a reference pulse in a nonlinear optical crystal [10,11]. The reference pulse is delayed by a variable time from the illuminating pulse, and the nonlinear process creates a high speed optical gate. If IsðtÞ is the signal and IrðtÞ is the reference pulse, the response function Sð Þ is given as

ð1

Nonlinear optical gating measures the time delay and intensity of a high speed optical signal. The time resolution is determined by the pulse duration, and the sensitivity is determined by the conversion efficiency of the nonlinear process. Optical time-of-flight ranging measurements were first demonstrated in biological tissues to measure corneal thickness and the depth of the stratum corneum and epidermis [12]. Dynamic ranges of 106 or higher can be achieved. Nonlinear cross correlation does not require pulses of as high intensity as the Kerr shutter but still requires the use of short pulses. It is also important to note that this technique detects the intensity (rather than the field) of the backscattered or backreflected light.

Interferometric detection overcomes many of the limitations of nonlinear gating techniques and can measure the echo time delay or backreflected or backscattered light with high dynamic range and high sensitivity. These techniques are analogous to coherent optical detection in optical communications (in contrast to direct detection). OCT is based on low coherence interferometry or white light interferometry, first described by Sir Isaac Newton. More recently, low coherence interferometry has been used to characterize optical echoes and backscattering in optical fibers and waveguide devices [13–15]. The first biological application of low coherence interferometry was in ophthalmologic biometry for measurement of eye length [16]. Since then, related versions of this technique have been developed for noninvasive high precision and high resolution biometry [17,18]. A dual beam interferometer was used to perform the first in vivo measurements of axial eye length in ophthalmology [19]. High resolution measurements of corneal thickness in vivo were demonstrated by using standard low coherence interferometry [20].

specimen, is lS, then the intensity of the interferometer output will oscillate as a function of l ¼ lR lS:

If the position of the reference mirror is varied, then the path length that the optical beam travels in the reference arm changes, and interference will occur. This is shown schematically in Fig. 2. If the light is highly coherent (narrow linewidth) or has a long coherence length, then interference oscillations will be observed for a wide range of relative path lengths of the reference and measurement arms. For applications in optical ranging or optical coherence tomography, it is necessary to measure absolute distance and dimensions of structures within the material or tissue. In this case, low coherence length light (broad bandwidth) is used. Low coherence light may be thought of as a superposition of electromagnetic fields with statistical discontinuities in phase as a function of time. The field is composed of different frequencies or wavelengths rather than a single wavelength. Low coherence light can be characterized by its coherence length (lc). The coherence length is inversely proportional to the bandwidth. When low coherence light is used in the interferometer, interference is observed only when the path lengths of the reference and measurement arms are matched to within the coherence length. This phenomenon is shown schematically in Fig. 2. The interferometer measures the field autocorrelation of the light. For the purposes of ranging and imaging, the coherence length of the light determines the resolution with which optical range or distance can be measured. The echo time delay and magnitude of reflected light can be determined by scanning the reference mirror position and demodulating the interference signal from the interferometer.

1.4RESOLUTION LIMITS OF OPTICAL COHERENCE TOMOGRAPHY

The next sections describe mechanisms that govern the performance of optical coherence tomography. Because OCT is based on modern optical communications technology, its performance can be predicted using well-established theories and engineered with high accuracy. Detailed specifications are not given here because specifications are dependent on instrument design and will change as different embodiments of this technology are developed.

In contrast to conventional microscopy, in OCT the mechanisms that govern the axial and transverse image resolution are independent. The axial resolution in OCT imaging is determined by the coherence length of the light source, and high axial resolution can be achieved independently of the beam-focusing conditions. The coherence length is the spatial width of the field autocorrelation produced by the interferometer. The envelope of the field autocorrelation is equivalent to the Fourier transform of the power spectrum. Thus, the width of the autocorrelation function, or the axial resolution, is inversely proportional to the width of the power spectrum. For a source with a Gaussian spectral distribution, the axial resolution z is

define two limiting cases for OCT imaging. Typically, OCT imaging is performed with low numerical aperture focusing to have a large depth of field, and low coherence interferometry is used to achieve axial resolution. In this limit the confocal parameter is larger than the coherence length, b > z. The axial image resolution is determined by the coherence length, and the transverse resolution by the spot size. In contrast to conventional microscopy, this mode of operation achieves high axial resolution independently of the available numerical aperture. This feature is particularly powerful for applications such as retinal imaging in ophthalmology or in catheter-endoscope-based imaging where the available numerical aperture may be limited. However, operation in the low numerical aperture limit yields low transverse resolution.

Conversely, it is also possible to focus with high numerical aperture and achieve high transverse resolution at the expense of reduced depth of focus. This operating regime is typical for conventional microscopy or confocal microscopy. Depending upon the coherence length of the light, the depth of field can be shorter than the coherence length, b < z. In this case the depth of field can be used to differentiate backscattered or backreflected signals from different depths. This regime of operation has been referred to as optical coherence microscopy (OCM). This mode of operation can be useful for imaging scattering systems because the coherence gating effect removes the contributions from scattering in front and in back of the focal plane more effectively than confocal gating. A comprehensive discussion of high transverse resolution OCT is provided in Chapter 10.

Although high transverse resolutions can be achieved in OCT, the depth of field will be limited. To perform imaging over a range of depths it is necessary to track the focal depth along with the axial range (reference delay) that is being detected [21]. Alternatively, it is possible to use a technique analogous to ultrasound C-mode scanning and acquire multiple images with different zones of focus and fuse these images together to create a single image with an extended depth of field [22].

1.5SENSITIVITY LIMITS FOR OPTICAL COHERENCE TOMOGRAPHY

Optical coherence tomography can achieve high detection sensitivity because interferometry measures the field rather than the intensity of light using optical heterodyne detection. This effect can be seen from Eq. (4), which describes the intensity of the interferometer output signal. The oscillating interference term is the result of the backscattered or backreflected electric field from the sample (which can be very weak) multiplied by the electric field of the reference beam. Because the beam from the reference mirror can have a large field amplitude, the weak electric field from the sample beam is multiplied by the large field, thereby increasing the magnitude of the oscillating term that is detected by the detector. The interferometer thus produces heterodyne gain for weak optical signals.

Backreflected or backscattered optical echoes from the specimen are detected by electronically demodulating the signal from the photodetector as the reference mirror is translated. In most OCT embodiments the reference mirror is scanned at a constant velocity v that Doppler shifts the reflected light. This modulates the interference signal at the Doppler beat frequency fD ¼ 2v= , where v is the reference mirror velocity. By electronically filtering the detected signal at this frequency, the

presence of echoes from different reflecting surfaces in the biological specimen may be detected. In addition, it is interesting to note that if the light is backreflected or backscattered from a moving structure, it will also be Doppler shifted and result in a shift of the beat frequency. This principle has been used to perform OCT measurements of Doppler flow.

The signal-to-noise performance can be calculated using well-established methods from optical communication. The signal-to-noise ratio (SNR) is given by the expression

SNR ¼ 10 logð P=2h NEBÞ ð8Þ

where is the detector quantum efficiency, h is the photon energy, P is the signal power, and NEB is the noise equivalent bandwidth of the electronic filter used to demodulate the signal. This expression implies that the signal-to-noise ratio scales as the detected power divided by the noise equivalent bandwidth of the detection. This means that high image acquisition speeds or higher image resolutions require higher optical powers for a given signal-to-noise ratio. The performance of optical coherence tomography systems varies widely according to their design and data acquisition speed requirements. However, for typical measurement parameters, sensitivities to reflected signals in the range of 90 to 100 dB can be achieved, corresponding to the detection of signals that are 10 9 or 10 10 of the incident optical power.

1.6IMAGE GENERATION AND DISPLAY IN OPTICAL COHERENCE TOMOGRAPHY

Optical coherence tomographic cross-sectional imaging is achieved by performing successive axial measurements of backreflected or backscattered light at different transverse positions [6]. Figure 4 shows one example of how optical coherence tomography is performed. A two-dimensional cross-sectional image is acquired by scanning the incident optical beam, performing successive rapid axial measurements of optical backscattering or backreflection profiles at different transverse positions. The result is a two-dimensional data set in which each trace represents the magnitude of reflection or backscattering of the optical beam as a function of depth in the tissue.

A wide range of OCT scan patterns are possible as shown in Fig. 5. The most common method of OCT scanning acquires data with depth priority. However, it is also possible to acquire data with transverse priority, by detecting backreflections or backscattering at a given depth or range while transversely scanning the imaging beam. A cross-sectional image can be generated by detecting the backscattering along successive x scans for different z depths. It is also possible to perform OCT imaging in an en face plane [23,24]. In this case the backreflected or backscattered signals are detected at a fixed z depth, scanning along successive x and y directions. This mode of imaging is analogous to that used in confocal microscopy. A comprehensive discussion of OCT imaging using other scanning protocols is presented in Chapter 11.

For purposes of visualization, OCT data are usually acquired by computer and displayed as a two-dimensional gray-scale or false color image. Figure 6 shows an example of a tomographic image of the retina displayed in gray scale. The image displayed consists of 200 pixels (horizontal) by 500 pixels (vertical). The vertical direction corresponds to the direction of the incident optical

Figure 4 Image generation in OCT. Cross-sectional images are constructed by performing axial measurements of the echo time delay and magnitude of backscattered or backreflected light at different transverse positions. This results in a two-dimensional data set that represents the backscattering in a cross-sectional plane of the material or tissue being imaged.

beam and the axial data sets. The optical beam was scanned in the transverse direction, and 200 axial measurements were performed. The backscattered signal ranges from approximately 50 dB, the maximum signal, to 100 dB, the sensitivity limit. Because the signal varies over five orders of magnitude, it is convenient to use the logarithm of the signal to display the image. This expands the dynamic range of the display but results in compression of relative variations in signal. In general the signal fluctuation in an OCT image is relatively high, so OCT has a limited capability for detecting small relative changes in backscattering, and it is difficult to detect relative changes in signal on the percent scale. In contrast, conventional imaging using charge-coupled devices (CCD) detectors can differentiate relative changes in signal of 10 4–10 6.

The intensity of the backscattered optical signal in an OCT image is typically represented on a logarithmic scale. The white level corresponds to the highest reflection or backscatter in the signal, and the black level corresponds to the weakest backreflection. The background noise in the image is typically thresholded and set to black. The gray-scale tomographic image differentiates structure in the retina including intraretinal layers and the retinal nerve fiber layer. The retinal pigment epithelium is highly scattering because it contains melanin. The choroid is highly vascular and is also highly scattering.

The dynamic range of gray-scale images is extremely limited. Most computer monitors provide only 8 bits or 256 gray levels. In addition, the eye has a limited ability to differentiate gray levels, so gray-scale images do not faithfully represent the full dynamic range of information available in OCT images. To enhance the differentiation of different structures within the image, the image may also be displayed in a false color representation as shown in Fig. 7. In this case, the logarithm of the