Optical coherence tomography (OCT), sometimes referred to as “optical biopsy”, can be used to obtain high-resolution (˜10 μm) cross-sectional imaging of scattering biological tissues up to 3 mm deep. OCT is based on low-coherence interferometery and fiber optic technology. The core of an OCT system is a Michelson interferometer. FIG. 1 shows the schematic of a simplified conventional fiberoptic OCT system 100. The OCT includes two optical fibers, shown as Fiber 1 and Fiber 2. Fiber 1 is used for the reference arm of the interferometer, while Fiber 2 is used for the sample arm of the interferometer. The reference arm (Fiber 1) is external to the probe, while the sample arm including Fiber 2 and the optical scanning mechanisms (e.g., transverse scanning mirror 130 (or rotating mirror)) are embedded inside the imaging probe for scanning the sample to be imaged, such as within a catheter for insertion into a body cavity of a patient.
Optical interference is detected by the photodetector 110 only when the optical path difference of the reference and sample arms is within the coherence length of the broadband light source 120. So, the depth (i.e., z-axis) information of the sample is acquired through the axial scanning (z) of a reference mirror in the reference arm. The lateral (i.e., x-axis) information is acquired through transversely scanning mirror 130. Therefore, two-dimensional (2D, i.e., x-z) cross-sectional images are obtained by transverse scanning mirror 130. 3D images can also be obtained if a 2D transversely x-y scanning mirror is used.
The axial resolution is determined by the coherence length of the light source. Low coherence is obtained by using a broadband light source such as a superluminescent diode (SLD) or a femtosecond laser. The coherence length of a broadband light source is given by 0.44λo2/Δλ, where λo and Δλ are respectively the center wavelength and spectral bandwidth of the light source. For example, a SLD with a center wavelength of 1300 mn and a bandwidth of 90 nm has a coherence length of 8 μm which is roughly the OCT axial resolution. Thus, OCT imaging can achieve at least one order of magnitude higher spatial resolution compared to commonly used ultrasound imaging (˜100 μm). Furthermore, study shows that more than 85% of all cancers originate in the epithelial layer which is within the penetration depth of infrared laser beams. Thus, OCT can be used for cancer diagnosis and has been applied to a wide variety of biological tissue and organ systems including eyes, skin, teeth, gastrointestinal tracts and respiratory tracts.
For intravascular applications such as in lung bronchi, gastrointestines and heart arteries circumferential (360°) scanning must be provided. Currently, there are several techniques used to provide circumferential (360°) scanning. One method involves rotating a long optical fiber with a prism at the fiber distal end. The rotating method is slow and has poor angular position control. Another method involves using a prism mounted on a micromotor. The micromotor method poses packaging difficulties and the fabrication of micromotors with less than about 2 mm in diameter. Moreover, having a large output torque is challenging. In yet another method, an imaging probe is moved back and forth several times with the imaging probe only covering a portion of the circumference. This method takes a much longer time, and increases the discomfort of the patient. The results of this method lack accuracy because of the motion of testing organs and the large time difference between adjacent scans.
What is needed is a fast scanning and the low cost miniature full circumferential scanning OCT probe that avoids either using expensive micromotors or slow rotation of optical fibers required by conventional OCTs to obtain full circumferential scanning. The size of the OCT probe should enable it to be disposed inside a conventional catheter.