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

because all fibers (and thus the tissue optic axis) are parallel to the incident light propagation direction.

9.5.6Conclusions

Depth-resolved Stokes parameter images of the corneas and retinas of New Zealand White rabbit eyes were acquired with PS-OCT. We demonstrated that the cornea is birefringent and that the birefringence as well as the optic axis of the stroma vary as a function of lateral position. PS-OCT could be used to study the corneal stroma structure more extensively.

We also imaged the retina and observed a birefringent layer at the vitreous/retina interface. The local thickness of the birefringent layer determined using PS-OCT correspond closely to the thickness of the NFL measured from histology slides in four out of six predetermined points. With SLP and OCT, NFL thickness is typically measured in a cylindrical ring around the optic nerve with a diameter of 1.5–1.75 optic discs, corresponding to the location of points 5 and 6 in our experiment. At these points, NFL thickness measured by PSOCT is in close agreement with the histological results. Furthermore, the magnitude of the birefringence is related to the axon density, and a quantitative analysis of the phase retardation could provide information on axon density within the NFL.

9.6FUTURE DIRECTIONS IN PS-OCT

The potential biological and medical applications of PS-OCT are just beginning to be explored. Much work remains for further development of PS-OCT. We anticipate that progress will proceed in three major areas: instrumentation, biological and medical applications, and data interpretation and image processing. Many clinical applications of PS-OCT will require the development of a fiber-based instrument that can record images at frame rates comparable to those of current OCT systems ( 5 fps). Because many components in biological materials contain intrinsic and/or form birefringence, PS-OCT is an attractive technique for providing an additional contrast mechanism that can be used to image and/or identify structural components. Moreover, because functional information in some biological systems is associated with transient changes in birefringence, the possibility of functional PS-OCT imaging should be explored. PS-OCT may hold considerable potential for monitoring, in real time, laser surgical procedures involving birefringent biological materials. Because many laser surgical procedures rely on a photothermal injury mechanism, birefringence changes in subsurface tissue components measured using PS-OCT may be used as a feedback signal to control dosimetry in real time. Finally, many features of PS-OCT interference fringe data require additional interpretation and study. Because polarization changes in light propagating in the sample may be used as an additional contrast mechanism, the relative contribution of light scattering and birefringence-induced changes requires further study and clarification. In principle, one would like to distinguish polarization changes due to scattering and birefringence at each position in the sample and use each as a potential contrast mechanism. In conclusion, we anticipate that PS-OCT will continue to advance rapidly and be applied to novel problems in clinical medicine and biological research.

ACKNOWLEDGMENTS

Research grants from the Institute of Arthritis, Musculoskeletal and Skin Diseases and Heart, Lung and Blood, and the National Center for Research Resources at the National Institutes of Health, U.S. Department of Energy, Office of Naval Research, the Whitaker Foundation, and the Beckman Laser Institute Endowment are gratefully acknowledged.

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systems used to date is sufficient to visualize normal morphology at the tissue level as well as disruptions of normal morphology due to disease processes. However, the resolution of current OCT systems is at least one order of magnitude less than that required to visualize most types of living cells; thus the goal of a true ‘‘optical biopsy’’ capable of conclusive diagnosis of pathologies at the cellular level (such as cancer) remains unrealized.

A number of approaches have been conceived to extend the spatial resolution of OCT. The most straightforward approach makes use of the well-known relation of the full width at half-maximum (FWHM) coherence length lc of a light source to its center wavelength and FWHM spectral bandwidth (under the assumption of a Gaussian source frequency spectrum) [1]:

By using shorter wavelength or broader bandwidth light sources for OCT, the coherence length of the source is thus reduced and the axial resolution is increased. Typical parameters for current clinical implementations of OCT include ¼ 840 nm,30 nm, z 10:4 m for ophthalmic OCT using superluminescent diode sources [2], and ¼ 1300 nm, 70 nm, lc 10:7 m for endoscopic OCT using semiconductor optical amplifier sources [3,4]. OCT systems using broader bandwidth sources including LEDs, broadband fluorescent sources, and modelocked lasers have also been reported featuring coherence lengths ranging as short as 2 m [5–7].

Optical coherence tomographic systems based on ultrabroad bandwidth sources have illustrated dramatic improvements in image quality, including a recent demonstration of cellular resolution imaging in Xenopus with 1:5 m 3 m (longitudinal transverse) resolution imaging using a novel sub-two optical cycle mode-locked Ti : sapphire laser [8]. However, significant signal losses and increased system complexity accompany the use of such sources, including the reduced heterodyne efficiency for shorter coherence length, difficulties in obtaining ultrabroadband fiber couplers, greatly increased difficulties of dispersion matching in the interferometer arms, and (in the case of femtosecond lasers) the dramatically increased complexity of the source itself. In addition, it is noteworthy that decreasing the source coherence length increases the resolution only in the axial dimension. Although the near-surface structures of many biological tissues accessible to OCT demonstrate a layered structure in which axial resolution is more critical than lateral resolution, most biological cell types do not exhibit this structure. Thus, a method for OCT imaging that increases resolution in all three spatial dimensions remains desirable.

A second approach for increasing the resolution of OCT springs from the observation that an OCT scanner is inherently a single spatial mode confocal microscope as well as a range-gating instrument [9,10]. Fundamentally, confocality of the light incident on and returning from the sample is ensured by the heterodyne mixing process, which selects out only the light returning from the sample that occupies the same spatial mode as the local oscillator (the reference light). In fact, OCT systems implemented in single-mode fiber optics are redundantly confocal, because the sample arm fiber also acts as a single-mode aperture for illumination and collection of light from the sample. In contrast to bright-field microscopy, in which the received power in the reflection geometry does not depend on the scatter

depth, in a confocal system the reflected power is a function of the scatterer location in all three dimensions [11]. Thus, in addition to the depth resolution that OCT derives from low coherence interferometry (hereafter referred to as the ‘‘coherence gate’’), any OCT system is also characterized by a depth and lateral response that is purely a function of the sample arm optics (hereafter referred to as the ‘‘confocal gate’’).

Abbe´’s rule for the lateral resolution x of a coherent microscope is given by

[12]

where the denominator reduces to twice the numerical aperture (NA) of the objective in a confocal microscope operating in the retroreflection configuration (see Fig. 1). Increasing the NA can thus increase the lateral resolution, but it also

beam waist of a Gaussian beam [13]. In conventional OCT, a relatively low numerical aperture objective (NA < 0:2) is typically used in the sample arm for two reasons. First, even a relatively low numerical aperture objective still delivers a lateral resolution x of 10–20 m, thus giving more or less square pixels when the depth resolution is set by the coherence length of a typical OCT light source. Second, the low numerical aperture objective also generates a long depth of focus (b greater than several hundred micrometers), within which the coherence gate may be axially scanned without much loss of light recoupling back into the confocal gate. This latter reason is particularly germane in the design of clinical diagnostic

Figure 1 Schematic illustration of sample arm focusing in optical coherence tomography (OCT) and optical coherence microscopy (OCM). In OCT, low numerical aperture focusing is employed to provide a long depth of focus to enable cross-sectional imaging using axial coherence gate scanning. In OCM, sample arm light is focused with a high numerical aperture objective to create a minimal focal volume in the sample. The low coherence interferometric coherence gate is overlapped with the focal volume to provide enhanced scatter rejection of out-of-focus light, while the sample is scanned in either the xy or xz planes to create en face or cross-sectional images, respectively. (From Ref. 53.)

devices, which must be used with living (and moving) patients. For example, an endoscopic OCT probe exiting the distal end of a several-meters-long endoscope deep inside the gastrointestinal tract of a patient cannot be positioned with much greater than millimeter accuracy in position next to the tissue to be imaged (at least without a dramatic redesign of the endoscope). In this case, a long depth of focus in the OCT sample arm optics enables reasonable tissue images to be obtained even if the physician has relatively poor control over the position of the sample arm optics themselves.

Increasing the numerical aperture of the sample arm optics increases both the lateral and axial resolutions of the confocal microscope and is thus an alternative method to increase the resolution in OCT. As shown in the following sections, overlapping the coherence gate of OCT at the same depth as the confocal gate of the confocal microscope increases the depth to which high resolution confocal images can be acquired in highly scattering media. In this book, we define optical coherence microscopy (OCM) as the implementation of OCT with high numerical aperture optics in the sample arm, accompanied by some means to coordinate the coherence gate position to the confocal gate position (Fig. 1). Optical coherence microscopes have been demonstrated using both single and array detection techniques; the former will be discussed in the present chapter, and the latter in Chapter 11.

10.3CONFOCAL MICROSCOPY IN SCATTERING MEDIA

In this section, we review the principles of confocal microscopy including fiberoptic implementations and imaging performance in scattering media. Confocal microscopy has been successful in imaging up to very high resolution (< 1 m) in relatively weakly scattering biological media and a few scattering depths into highly scattering media [12,14–17]. Combining focused illumination with spatially filtered detection, the confocal microscope collects signal from a diffraction-limited focal volume and rejects background that originates outside that focal volume. In an ideal case using single-point detection, the image intensity of the microscope is given by [12]

where hill and hdet are the illumination and detection point spread functions (PSFs), respectively (given by the Fourier transforms of the condenser and objective pupil functions), u and are axial and lateral optical units, respectively (in terms of the actual axial and lateral distances z and x), and R is the power reflectivity of the

sample object. The quantity hconfocalðu; Þ is the product of the illumination and detection PSFs and represents the overall PSF of the confocal imaging system. In

the reflection mode of a confocal microscope, the same objective is used for sample illumination and light collection, and thus hill ¼ hdet. For an ideal point object, the power reflectivity in Eq. (3) simplifies to a three-dimensional delta function, and the axial and lateral components of the response function separate, given by

decrease as NA increases. Thus, the axial and lateral resolutions of a confocal microscope are improved at high numerical aperture.

The axial response of a confocal microscope is easier to measure experimentally by measuring the PSF of a uniform planar object such as a mirror by scanning the mirror axially through the focus. In this case, the reflectivity Rðu; Þ in Eq. (3) is equal to 1 at u ¼ 0 for all . The axial component of the image intensity is then modified to

For an objective with NA ¼ 0:4, the FWHM of the effective PSF, or the FWHM of the measured image intensity of an ideal mirror as a function of the depth, is approximately 8 m. For NA ¼ 0:3, this FWHM increases to 35 m.

Considering a finite-sized detector (or equivalently a finite-sized confocal pin-

hole) Dð Þ, Eqs. (4) and (5) are modified to

ð

where the signal intensity is integrated over the detection aperture. The axial discrimination or ‘‘confocality’’ is dependent on the size of the confocal pinhole. It has been shown that the confocality, using a pinhole radius vp in optical units of less than 2.5, is the same as that using an ideal point detector [12]. The optical unit is then defined in Eq. (3) as ð d=Þ sinðNAÞ, where d is the diameter of the confocal pinhole. The actual pinhole size is obtained by dividing vp by the system magnification M between the objective and the detector and should be less than 2:5M=½ sinðNAÞ& in an ideal point detection case.

The penetration depth of a confocal microscope is limited by the ratio of signal S to background B ðS=BÞ. The background noise originates from outside the focal volume if the pinhole is not infinitesimally small (neglecting multiple scattering; see below). The background noise can be greatly attenuated by decreasing the size of the confocal pinhole [14–17]. Assuming a uniform background, Webb et al. [15] and Sheppard et al. [16] demonstrated independently that the background rejection improves the signal-to-noise ratio. The ability of optical sectioning and image contrast is thus improved such that the signal-to-noise ratio rather than the signal-to- background ratio limit the information capacity of an image [14–17]. The penetration depth can be expressed in terms of mean free path (MFP), which is the product of the depth and the extinction coefficient t ¼ s þ a of the sample medium. In biological samples, the absorption coefficient a is usually negligible in the nearinfrared spectral region; thus the extinction coefficient is dominated by the scattering coefficient s. The penetration depth limit of a conventional confocal microscope has been estimated to be dependent on the numerical aperture and to be 5–9 MFP

using single backscatter theory (discussed in the next section) [9,18,19]. However, the penetration depth limit is degraded to 3–4 MFP if multiple scattering is considered [20]. The penetration depth also depends upon the reflectivity of the sample; a highly reflective grating has been resolved through 6 MFP of scattering media by using a confocal pinhole with diameter equal to 1.3 optical units [21].

The single-mode fiber-optic confocal microscope is a relatively new implementation of the confocal scanning microscope. The introduction of optical fibers makes the system more compact with numerous potential remote imaging applications, but it also changes the nature of the imaging point spread function. Because the light source in a fiber confocal microscope is the spatial mode profile of the light exiting the fiber tip (which also defines the confocal pinhole), image formation in this system is different from that of a conventional confocal microscope. In this case, the axial effective point spread function of a planar reflective sample from Eqs. (4) is modified to [22]

Here, the system magnification is assumed to be 1. The dimensionless parameter A denotes the normalized fiber spot size, where a0 is the pupil radius of the objective, r0 is the radius of the core of a single-mode fiber, and d is the distance from the fiber tip to the collimating lens. Parameter A plays a role analogous to that of the normalized pinhole radius p in a hard-aperture confocal microscope. For an objective with a fixed NA, the width of the PSF becomes broader as A increases. Thus, the strength

of the optical sectioning effect (or degree of confocality) decreases with increasing A. p

d ¼ 4:5 mm, a0 mm, and r0 ¼ 4:5 um, A is less than 0.5. Since hconfocal;planarðuÞ for A ¼ 0 and A ¼ 1 are nearly identical, the degree of confocality for A < 1 is close to

the case of an ideal single-point detection in a confocal microscope. For the case of a point reflector and in the limit A ¼ 0, the axial PSF of the single-mode fiber confocal microscope simplifies to be the same as the axial component of Eqs. (4) [22].

Unlike in a conventional confocal microscope, however, the received intensity in a fiber-optic confocal microscope is limited by the nonadjustable fiber core size. In the conventional microscope, the confocal pinhole can be opened up to 2:5vp to have image quality equivalent to using an ideal point detector but with higher signal intensity. This flexibility is not available in the optical fiber implementation.

Although fiber-optic confocal microscopes have strong potential for use in in vivo medical diagnostics, most confocal microscopes available commercially for general-purpose applications are table-top systems using bulk optics. This implementation provides the advantages of flexibility in the choice of light sources, scan parameters, and pinhole size. The greatest impediment to the widespread use of fiber-optic confocal microscopes in remote imaging applications such as endoscopy