Surgery – Diagnostic testing – Detecting nuclear – electromagnetic – or ultrasonic radiation
Reexamination Certificate
1999-03-19
2002-04-09
Lateef, Marvin M. (Department: 3737)
Surgery
Diagnostic testing
Detecting nuclear, electromagnetic, or ultrasonic radiation
C600S182000, C606S010000
Reexamination Certificate
active
06370422
ABSTRACT:
BACKGROUND OF THE INVENTION
1. Field of the Invention
The present invention relates generally to the fields of optics and microscopy. More particularly, it concerns apparatus and methods for analyzing samples using fiber-optic confocal imaging techniques.
2. Description of Related Art
Currently, two methods are available for imaging at the cellular level in the body: Optical Coherence Tomography (OCT) (Izatt et al., 1994; Huang et al., 1991; Schmitt et al., 1993; Yadlowsky et al., 1995; Swanson et al., 1993; Schmitt et al., 1994) and confocal imaging (Petroll et al., 1993; Jester et al., 1991; Rajadhyaksha et al., 1995a; Gmitro and Aziz, 1993; Massig et al., 1994; Masters and Thaer, 1994; Giniunas et al., 1993a; 1993b; Delaney et al., 1994).
OCT uses interference techniques with low-coherence light sources to select the light coming from a distinct depth (Huang et al., 1991). The basic system uses a Michaelson interferometer with the tissue sample in one arm and a reference mirror in another arm. When the reflections from both arms are combined at the detector, an interference maximum or minimum is detected when the reflections from both arms are matched in optical path length (time-of-flight). The strength of the interference is proportional to the amount of light reflected from the corresponding optical path length within the tissue. The temporal frequency of the interference maximum and minimum can be modulated by translating the reference mirror at a constant velocity or by stretching the path length in the reference arm with a piezo-electric transducer at the modulation frequency. These systems use heterodyne detection of the modulated interference signal to detect as little as 5×10
−10
of the incident light (Huang et al., 1991). The axial resolution of OCT depends on the coherence length of the illumination source, which is in the range of 10 to 20 &mgr;m (Swanson et al., 1993) for semi-conductor sources. More recent work with mode-locked Ti:Sapphire and Forrestrite lasers have yielded coherence lengths as small as 1.8 &mgr;m (SPIE Proceedings, 1997). The lateral resolution is determined by the diffraction-limited spot size in the tissue.
OCT forms a cross-sectional image of the tissue by mapping the intensity of the reflected light as the sampling point is translated in the axial and lateral dimensions. The sampling point is moved in depth by the translation of the reference mirror. The lateral dimension is achieved by translation of the optics over the surface of the tissue. Much of the in vivo imaging done with OCT has been in the eye (Swanson et al., 1993). Some researchers have also attempted to use the technique to image scattering tissue such as human skin, however, have not obtained images of individual cells due to the lack of spatial resolution (Yadlowsky et al., 1995). Thus, while OCT has a high sensitivity, it has not demonstrated the spatial resolution necessary to image cellular structure. Although the new mode-locked laser gives OCT the potential to achieve the desired resolution, the cost and complexity of these lasers make them impractical for clinical use.
The first attempts at in vivo confocal imaging have been done with a modified scanning Nipow disk microscope (Petroll et al., 1993; Jester et al., 1991). A Nipow disk refers to fiat disk which has a staggered array of pinhole apertures spread over the entire disk. At one instant in time, one of the apertures passes the illumination light and detects the reflected confocal light. As the disk spins, the aperture being used for illumination and detection moves around the disk, thereby, imaging the entire sample. The disks can spin at very high speeds to produce images at video rates. These systems have been used to image in vivo cornea and several organs of a rat which had been exposed by laparotomy. The best spatial resolution reported to date is approximately 7 &mgr;m, with no mention of the sensitivity or corresponding maximum penetration depth. One limitation of this apparatus is the fact that these microscopes are susceptible to misalignment of the disk.
More recently, a confocal system has been developed with a spatial resolution sufficient to image individual skin cells in a living human (Rajadhyaksha et al., 1995b). The instrument is a simplified confocal microscope with a standard pinhole aperture and scanning mirrors. A high numerical apertune (NA) oil-immersion objective lens is used in contact with the skin to achieve a lateral resolution of approximately 2 &mgr;m. No values have been reported for the sensitivity of the system, however it is capable of imaging the entire thickness of the forearm epithelium and into the rete ridges of the underlying stroma using 830 nm light. Images of cell size and nuclear to cytoplasmic ratio obtained with this system agree well with those measured from biopsies, validating the concept that in vivo confocal imaging can be used to assess tissue morphology. However, the size and configuration of illumination optics prevents use of this system to image tissues within moderately accessible cavities such as the cervix or mouth.
Several authors have proposed fiber optic systems for in vivo confocal imaging (Massig et al., 1994; Masters and Thaer, 1994; Giniunas et al., 1993a; Giniunas et al., 1993b; Delaney et al., 1994) based upon fiber optics. These designs implement confocal detection through a single fiber optic. These designs also incorporate some method of translating the endpiece optics in the axial and transverse directions to form an image. Designing an endoscopic system encompassing a miniature, high speed mechanical scanning system with high spatial resolution is difficult.
Another approach to a fiber optic design (Gmitro and Aziz, 1993) uses a fiber optic imaging bundle as a confocal image conduit between the endpiece optics and a confocal microscope. The function of the confocal microscope was to scan the illumination spot across the fiber bundle and to detect the emerging light. This arrangement does avoid the need for a mechanical translation system in the endpiece, however a commercial confocal microscope is expensive and cumbersome, limiting its usefulness as a practical clinical tool. In addition, the high absorption and scattering of the tissue will not allow the fluorescence excitation and emission light to penetrate the entire depth of the epithelium.
More recently, the same design has been implemented for reflection imaging using white light (Juskaitis et al., 1997) Reflection imaging is capable of penetrating to greater depths; however, the use of a white light source will limit the illumination power available to the system. A consequence of the limited illumination power will be a limited penetration depth due to loss of signal in scattering tissue. The rationale given for using white light is eliminating the speckle observed when imaging a resolution test target with laser light.
Although the potential of both in vivo confocal imaging with subcellular resolution and of fiber optic confocal imaging has been demonstrated, there is currently not a system which provides the imaging capabilities required for imaging tissues in vivo within the physical constraints necessary to achieve sufficient resolution and magnification of tissues within a living organism.
The system described by Juskaitis et al. (1997) uses a combination of angle polishing and index matching with glycerin to “prevent” specular reflection from the faces of the fiber bundle. Because the fiber bundle requires a difference in index between the individual fiber cores and cladding to function, it is not possible to eliminate or “prevent” the reflection from the fiber face at the proximal end (i.e. the end that light is injected into). Instead, the reflections can be minimized by using a matching oil with an index half way between the index of the cores and the cladding. Indeed, Juskaitis et al. use such an oil. While it is theoretically possible to eliminate the reflections from the distal end by using an oil with an index exactly equal to the index of the fiber co
Bowman Brett S.
Descour Michael R.
Richards-Kortum Rebecca
Smithpeter Colin L.
Board of Regents , The University of Texas System
Fulbright & Jaworski
Lateef Marvin M.
Shaw Shawna J.
LandOfFree
Fiber-optic confocal imaging apparatus and methods of use does not yet have a rating. At this time, there are no reviews or comments for this patent.
If you have personal experience with Fiber-optic confocal imaging apparatus and methods of use, we encourage you to share that experience with our LandOfFree.com community. Your opinion is very important and Fiber-optic confocal imaging apparatus and methods of use will most certainly appreciate the feedback.
Profile ID: LFUS-PAI-O-2932806