Ординатура / Офтальмология / Английские материалы / Handbook of Optical Coherence Tomography_Bouma, Tearney_2002
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between resolution and working distance. In a high numerical aperture system the confocal parameter may be on the single micrometer scale; thus much more mechanical stability is required for the sample arm optics than in OCT.
As in conventional OCT, two-dimensional OCM images may be acquired in either the xz or xy planes (with reference to Fig. 1) or in some tilted plane in between. When working with a confocal gate of approximately the same dimension as the coherence gate, however, it is a challenging concern to maintain perfect overlap of the two gates while scanning. This is because the axial positions of the confocal and coherence gates have different dependences on the refractive index of the sample. Under the paraxial approximation, when a collimated beam is focused through an objective of focal length f into a medium having refractive index n, axial movement of the objective by a distance z results in a movement of the focal point by n z. The equivalent optical pathlength is thus moved by n2 z. Thus, any axial scan of the OCM objective must be accompanied by a coordinated reference arm scan of n2 times that amount [9,34]. At sufficiently high numerical aperture that the paraxial approximation cannot be used, this relationship holds true if index-matching fluid is used between the objective and the sample, but it becomes much more complicated if there is an air gap between the objective and the sample [9]. Focus-tracking scan coordination techniques based on deterministic rules may, of course, be achieved using computer-controlled translation stages. However, in samples with unpredictable variations in refractive index such as biological tissue, large variations in the index may decouple the confocal and coherence gates unless alternative means are used for their determination [48].
A second challenge associated with adding coherence gating to confocal microscopy is to implement phase modulation techniques for noise reduction in frequencyselective detection, particularly in en face imaging systems in which there is no Doppler shift from a moving reference arm. Several novel approaches have been reported for this purpose.
10.6.1En Face Imaging Systems
The simplest approach to deal with the focus-tracking problem in OCM is to avoid it by acquiring en face images in the xy plane after carefully coordinating confocal and coherence gates. This is the approach adopted by Izatt et al. [9,25,52], Podoleanu et al. [30–32,49], and Haskell et al. [50].
Izatt et al. [25] employed a straightforward approach in which a moderate numerical aperture objective in the sample arm (NA ¼ 0:65) was rigidly fixed to an optical table and the sample was raster scanned underneath. Reference arm phase modulations for lock-in detection was achieved by implementing small-ampli- tude ð< 1 mÞ modulation of the reference mirror position using a PZT stack. This system employed a method suggested by Chinn [51] to eliminate modulation phase instability due to random fiber length changes. For an arm-length-matched OCM system with a PZT-controlled reference mirror, the interferometric component of the detector current may be written in the form
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Vin |
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id ¼ 1 þ cos |
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V |
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where Vin is the piezo modulation voltage, v is the voltage required to produce a phase shift in the interferogram, and is a random phase shift due to fiber length variations. If the PZT voltage is modulated with the waveform Vin ¼ V sin t, then it can be shown that the detector will detect currents modulated at the first and second harmonics of the modulation frequency according to
i ¼ 2J1ð Þ sin ’ sin t; |
i2 ¼ 2J1ð Þ cos ’ cos 2 t |
ð26Þ |
The sum of the powers in the first and second harmonics, |
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i2 þ i22 ¼ J12ð Þ sin2 þ J22ð Þ sin2 |
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can then clearly be made independent of the random phase for the proper choice of the modulation amplitude such that J1ð Þ ¼ J2ð Þ, which occurs when ¼ 2:6. Thus, the interferometric fringe amplitude was measured invariantly with respect to random arm length drifts by monitoring the sum of the detector powers demodulated at both harmonics of the PZT driving frequency.
The Izatt system was used to image features of human colon mucosal tissue in vitro. Sample images are reproduced in Fig. 8 and remain among the only images recorded to date using a coherence technique in which individual cells are visualized hundreds of micrometers deep in a highly scattering biological tissue. More recently,
Figure 8 Optical coherence microscopic images corresponding to thin (5 m FWHM) optical sections obtained parallel to the mucosal surface at the depths indicated in a normal human mucosal specimen in vitro. Mucosal substructure details visualized using OCM include colonic crypt lumens (lu), simple columnar epithelial cells (ec), and lamina propria connective tissue (lp). An individual goblet cell (gc; arrow) is identified as containing a nonscattering spherical mucin droplet. The proportional cross section occupied by connective tissue (lamina propria) versus crypt structures increases near the crypt bases in the deep mucosa. The maximum OCM sectioning depth was 600 mm, limited by the available low coherence power. Bar length ¼ 200 m. (From Ref. 25.)
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the same group reported a rapid raster scan OCM system operating at several frames (4–8) per second [52].
Podoleanu et al. [30–32,49] used a novel sampling function to decode en face OCT images of the retina. A schematic diagram of their system is shown in Fig. 9, including two perpendicular galvanometric mirrors to form 2-D raster scans. The centers of both scanners were located on the optical axis of the sample arm optics. When one scanner was fixed, the optical pathlength changed when the other scanner swept the beam from the center of the image plane, O00, to N. The reference arm length was fixed to equal the sample pathlength at position O 00. The interferometric signal had maximum intensity when the optical path difference was a multiple of a wavelength. By placing a mirror in plane S and scanning two scanners, nonuniform concentric rings similar to Newton rings were observed. The rings were nonuniform
due to nonlinearity of the optical path differences between O00 |
and N. Since the |
light source was quasi-monochromatic, the rings appeared |
within an area |
A ð =2Þlcr assuming the optical path difference was within lc=4. Because the rings contained information about the scanning angles (as a function of time as well as frequency) for each position, the rings were used as a sample function such that the returned signal was modulated in intensity at the frequency f given by
f ¼ 8 |
x2 1 |
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kUx ! |
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where k is the scanner parameter in degrees per volt, and U and ! are the amplitude and frequency of the driven scanner, respectively.
Using a setup similar to that of Fig 9 but shifting the center of the scanners by, the sampling function changed from concentric rings to high density uniform lines [31,32]. Higher spatial frequency of the sampling function corresponds to higher modulation frequency, which reduced the contribution of 1=f noise. Using !x ¼ 300 Hz, !y ¼ 0:5 Hz, and ¼ 3 mm, in vivo images of the human retina were obtained. The spatial resolution was 6 m. The image acquisition rate was up to several frames per second.
Figure 9 Schematic of OCM using a Newton rings sampling function. (From Ref. 30.)
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10.6.2 Focus-Tracking Systems
An ingenious and elegant solution for axial focus tracking at relatively low numerical aperture that depends on the fact that in biological tissue n2 2 was demonstrated by Schmitt et al. [33,34]. In this implementation (Fig. 10), a retroreflecting prism in the reference arm and the objective lens in the sample arm were both mounted on the same translation stage. When the objective scanned a distance z toward the sample, the optical pathlength in the reference arm increased by 2 z to achieve the pathlength matching. An example of the excellent quality of images obtained in skin using this system is provided in Fig. 11.
Another focus-tracking technique that is suitable for high speed imaging was reported by Lexer et al. [35]. In this technique, a sample beam focus reflected offset from the center of a scanning galvanometer mirror was imaged into the sample. The imaging system translated the angular deviation of the scanner into a depth scan in the sample, and the index of the sample was compensated for by the proper choice of imaging system magnification. This technique was demonstrated for moderate speed (1 image/s) imaging of cellular structure in human corneal samples with an imaging resolution of 14 m 5 m (longitudinal transverse).
10.7CONCLUSION
Optical coherence microscopy is an extension of optical coherence tomography that capitalizes on the confocal nature of OCT to improve image resolution by imaging with high numerical aperture. From the point of view of practitioners of confocal microscopy, OCM adds the possibility of imaging deeper into highly scattering media at the cost of limiting both the quantity (due to single-mode detection) and nature (coherent scattering processes only, i.e., backscattering, are detected) of the detected light. From the point of view of practitioners of OCT, OCM is the only technique to increase spatial resolution in all three spatial dimensions: however, this comes at the cost of increased overall complexity of the imaging system. Much of the added complexity arises from the need to coordinate sample and reference arm scans
Figure 10 Configuration of optical coherence tomography using a dynamic focusing technique. (From Ref. 33.)
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Figure 11 OCM image of living human nail fold region using the dynamic focusing technique. (From Ref. 33.)
in OCM as well as to incorporate high quality optics in the sample arm. Although promising initial results have been obtained, the future of this technique depends upon several unanswered questions. As Fig. 6 illustrates, it is not yet clear over what parameter range OCM enhances confocal microscopy, although some clear improvements have been demonstrated. The potential of OCM in the medical environment rests primarily on the development of microoptical endoscopic systems capable of incorporating confocal-quality scanning optics into an endoscope or other diagnostic probe tip.
ACKNOWLEDGMENTS
We acknowledge research support from the National Science Foundation (BES9624617), the Whitaker Foundation, and Olympus Graduate Research Fellowship (H-W.W.).
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11
Full-Field Optical Coherence
Microscopy
H. SAINT-JALMES and M. LEBEC
Universite´ Claude Bernard-Lyon I, Villeurbanne, France
E. BEAUREPAIRE, A. DUBOIS, and A. C. BOCCARA
CNRS, ESPCI, Paris, France
11.1INTRODUCTION
With the development of new optical technologies and powerful digital image processors, optical microscopy has been moving beyond traditional two-dimensional (2- D) imaging to the reproduction of three-dimensional (3-D) objects. The demonstration in the late 1950s of the confocal scanning principle, with its capacity to reject light from out-of-focus planes, provided a powerful tool for reconstruction of 3-D objects [1]. Confocal laser scanning microscopy, introduced around 1980, is nowadays a well-established tool for biological and biomedical imaging [1–3]. More recently, optical coherence tomography (OCT) [4] was introduced as a novel high resolution noninvasive imaging technique best suited for imaging in scattering media [5]. OCT uses low coherence interferometry to reject scattered light and amplify the single-backscattered component emerging from the sample, and was demonstrated to be an effective technique to produce images with 10–15 m resolution several hundred micrometers deep inside various biological tissues [6–9]. Optical coherence microscopy (OCM), a variant of OCT that takes advantage of a high numerical aperture (NA) objective lens and produces XY (head-on) rather than XZ images, can provide an order-of-magnitude higher resolution [10–12]. In addition, coherent detection gives access to the optical phase, which can be exploited to image transparent objects, to measure birefringence, or to reconstruct the topography of surfaces.
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Full-field optical coherence microscopy [13] is an alternative method developed to retain advantages of scanning OCM systems while enabling simultaneous acquisition of all the pixels of an image. This technique borrows its basic principle from OCT: A low temporal coherence light source is used in a Michelson interference microscope to select single-backscattered photons coming from a given depth in the sample. The main difference is that full-field systems avoid any X-Y scanning by illuminating the whole field of view with a spatially incoherent light source and by taking advantage of a detector array and a parallel coherent detection scheme. A complete slice orthogonal to the Z axis is thus acquired in one shot (Fig. 1). The only mechanical scanning involved in a complete 3-D image acquisition is a relatively slow motion along the Z axis in order to obtain images from different depths in the sample.
Section 11.2 presents the concepts involved in full-field coherent imaging and outlines some advantages of the parallel detection and head-on geometry in the
Figure 1 XY imaging. In a single-detector experiment, three mechanical movements are required to scan the volume of the sample. With an array of detectors, a plane perpendicular to the Z axis is acquired simultaneously; only a slow Z translation is necessary to acquire the 3-D image of the sample.