Patent Publication Number: US-11647905-B2

Title: Optical coherence tomography with graded index fiber for biological imaging

Description:
CROSS REFERENCE TO RELATED APPLICATIONS 
     This application is a continuation of U.S. patent application Ser. No. 15/419,815, titled “OPTICAL COHERENCE TOMOGRAPHY WITH GRADED INDEX FIBER FOR BIOLOGICAL IMAGING,” filed Jan. 30, 2017, now U.S. Pat. No. 10,952,615, which is a continuation of U.S. patent application Ser. No. 14/400,140, titled “OPTICAL COHERENCE TOMOGRAPHY WITH GRADED INDEX FIBER FOR BIOLOGICAL IMAGING”, filed Nov. 10, 2014, now U.S. Pat. No. 9,557,156, which is a 35 U.S.C. § 371 national phase application of International Patent Application No. PCT/US2013/031951, titled “OPTICAL COHERENCE TOMOGRAPHY WITH GRADED INDEX FIBER FOR BIOLOGICAL IMAGING”, filed Mar. 15, 2013, Publication No. WO 2013/172972, which claims priority to U.S. Provisional Patent Application No. 61/646,783, titled “OPTICAL COHERENCE TOMOGRAPHY WITH GRADED INDEX FIBER FOR BIOLOGICAL IMAGING”, filed May 14, 2012, each of which is incorporated by reference in its entirety. 
     This application may be related to U.S. patent application Ser. No. 12/790,703, titled “OPTICAL COHERENCE TOMOGRAPHY FOR BIOLOGICAL IMAGING”, filed May 28, 2010, Publication No. US-2010-0305452-A1, which is incorporated by reference in its entirety. 
    
    
     INCORPORATION BY REFERENCE 
     All publications and patent applications mentioned in this specification are herein incorporated by reference in their entirety to the same extent as if each individual publication or patent application was specifically and individually indicated to be incorporated by reference. 
     FIELD 
     Described herein are imaging devices and systems for use in biological probes. In particular, described herein are catheter-based imaging systems using Optical Coherence Tomography (OCT). 
     BACKGROUND 
     In intravascular surgery, as well as other medical applications, there is frequently a need to extend very thin (few millimeter diameter), long (30-150+cm), and sterile catheters into thin-walled (e.g., 1-1.5 millimeter wall thickness) biological lumens, including blood vessels such as arteries and veins. 
     A number of vascular diseases, such as coronary artery disease and peripheral vascular disease, are caused by the build-up of atherosclerotic deposits (plaque) in the arteries, which limit blood flow to the tissues that are supplied by that particular artery. Disorders caused by occluded body vessels, including coronary artery disease (CAD) and peripheral artery disease (PAD) may be debilitating and life-threatening. Chronic Total Occlusion (CTO) can result in limb gangrene, requiring amputation, and may lead to other complications and eventually death. Increasingly, treatment of such blockages may include interventional procedures in which a guidewire is inserted through a catheter into the diseased artery and threaded to the blocked region. There the blockage may be either expanded into a more open position, for example, by pressure from an inflated catheter balloon (e.g., balloon angioplasty), and/or the blocked region may be held open by a stent. Treatment of such blockages can also include using a catheter to surgically remove the plaque from the inside of the artery (e.g., an atherectomy). 
     There is medical interest in equipping catheters with sensors that can help direct the catheter for atherectomy, occlusion-crossing, and/or other surgical procedures. For example, it would be useful to have sensors that can give the surgeon immediate visual feedback as to whether a particular tissue is diseased and/or how far away the cutting portion of a catheter is from the boundary of a particular blood vessel layer to minimize the risk of accidental damage. Conventional radiological imaging methods and ultrasound imaging systems have been attempted for such surgical procedures. However, neither ultrasound nor radiological imaging methods have enough resolution to help guide the operation of the catheter through small dimensions. Moreover, standard radiological techniques cannot easily discriminate between healthy tissue and diseased tissue unless the tissue has become heavily calcified. Further, the components of an ultrasound system are generally too large to implement on a small scale, such as with a system configured to be used within blood vessels. 
     Optical Coherence Tomography (OCT) has been proposed as one technique that may be particularly helpful for imaging regions of tissue, including within a body lumen such as a blood vessel. At a basic level, OCT relies on the fact that light traveling from a source and scattering from more distant objects takes longer to travel back than light scattering from nearby objects. Due to the wave nature of light, very small timing differences caused by light signals traveling different distances on the micron scale can cause constructive or destructive interference with reference light signals. OCT systems measure the resulting interference to obtain an image of the target. A typical OCT system requires one or more interferometers to distinguish the signal from the applied light. In addition, most known OCT systems, when applied to catheters, include a fiber that is rotated (often at high rates) within the catheter in order to scan the lumen and a second, large reference arm. 
     Referring to  FIG.  1   , a typical OCT device includes a target arm and a reference arm to generate a reference signal. In order to provide the interference reference signal, the OCT device will split an illuminating light signal from the source in two equal or unequal parts, send part of the illuminating light to the target of interest through one target optical “target arm” and send the other part of the illuminating light down a separate reference arm. Light from the separate reference arm reflects off of a mirror, and then returns and interferes with the scattered light that is returning from the target optical arm after bouncing off of the target. In a traditional OCT device, the reference arm length is engineered to be exactly the same length as the target arm so that the interference effect is maximized. The resulting interference between the two beams creates interference that can be measured to extract depth information related to the target. Using this depth information, an image of the object can be generated. Referring still to  FIG.  1   , a typical OCT device can further include a focusing lens in the target arm, such as a graded index (GRIN) lens, configured to focus the light coming out of the optical fiber into the tissue. 
     These traditional OCT systems, however, are large and cumbersome due to the required reference arm and are therefore generally ineffective for use in a medical catheter, particularly for use with a low cost and disposable catheter. Using a common path OCT system, i.e., a system without a separate reference arm, is one way to eliminate the cost and size of such an imaging catheter. There are several challenges, however, associated with developing a catheter having common path OCT. For example, a common path OCT system requires that the reference reflection be formed within the same optical conduit as the target reflection. This reference reflection must be finely tuned to avoid noise in the system, requiring that the path from the light source to the reflection interface be free of unnecessary components, such as focusing elements, that could interfere with the reference reflection. Further, the common path system must have components that are small enough to fit inside of a single small catheter, making it difficult to include additional components. Finally, for common path OCT, it is desirable to have the reference reflection as close to the tissue as possible to maintain the imaging range within the coherence length of the source and avoid data processing burden, as data processed for the distance between the reference and the start of the imaging is not useful. Accordingly, a common path OCT system that solves some of these problems is desired. 
     SUMMARY OF THE DISCLOSURE 
     In general, in one embodiment, a system for optical coherence tomography includes a source of optical radiation, an optical fiber, and a graded index fiber attached to a distal end of the optical fiber. The optical fiber and the graded index fiber are together configured to provide a common path for optical radiation reflected from a reference interface at a distal end of the graded index fiber and from a target. The system further includes receiving electronics configured to receive the optical radiation reflected from the reference interface and the target and a processor to generate an image of the target based upon the optical radiation received by the receiving electronics. 
     This and other embodiments can include one or more of the following features. A secondary reflection of optical radiation from an interface between the optical fiber and the graded index fiber can be less than −60 dB. The reference interface can provide a reference reflection of between −28 and −42 dB. The system can further include an interface medium at the reference interface that can be in optical contact with the graded index fiber. A refractive index of the graded index fiber and a refractive index of the reference medium can be mismatched such that the receiving electronics operate in a total noise range that can be within 5 dB of the shot noise limit. The reference interface can be angled with respect to a longitudinal axis of the graded index fiber such the receiving electronics can operate in a total noise range that is within 5 dB of the shot noise limit. A surface of the interface medium that is closest to the target can be concave. The interface medium can be an adhesive. An outer diameter of the graded index fiber and a protective coating around the graded index fiber can be less than 0.01 inches. 
     In general, in one embodiment, a system for optical coherence tomography includes a source of optical radiation, an optical fiber, and a graded index fiber attached to a distal end of the optical fiber. The optical fiber and the graded index fiber are together configured to provide a common path for optical radiation reflected from a reference interface at a distal end of the graded index fiber and from a target. The system further includes receiving electronics configured to receive the optical radiation reflected from the reference interface and the target and a processor to generate an image of the target based upon the optical radiation received by the receiving electronics. A secondary reflection of optical radiation from an interface between the optical fiber and the graded index fiber is less than −60 dB. 
     This and other embodiments can include one or more of the following features. The reference interface can provide a reference reflection of between −28 and −42 dB. The system can further include an interface medium at the reference interface that can be in optical contact with the graded index fiber. A refractive index of the graded index fiber and a refractive index of the reference medium can be mismatched such that the receiving electronics operate in a total noise range that can be within 5 dB of the shot noise limit. The reference interface can be angled with respect to a longitudinal axis of the graded index fiber such the receiving electronics can operate in a total noise range that is within 5 dB of the shot noise limit. A surface of the interface medium that is closest to the target can be concave. The interface medium can be an adhesive. An outer diameter of the graded index fiber and a protective coating around the graded index fiber can be less than 0.01 inches. 
     In general, in one embodiment, a catheter for use with optical coherence tomography includes an elongate catheter body. The catheter includes an optical fiber in the elongate catheter body and a graded index fiber attached to a distal end of the optical fiber. The optical fiber and the graded index fiber are together configured to provide a common path for optical radiation reflected from a reference interface at a distal end of the graded index fiber and a target. A secondary reflection of optical radiation from an interface between the optical fiber and the graded index fiber can be less than −60 dB. 
     This and other embodiments can include one or more of the following features. The reference interface can provide a reference reflection of between −28 and −42 dB. A distance from an outer edge of the elongate catheter body and a focal point of the GRIN lens can be less than 1 mm. A distance from an outer edge of the elongate catheter body and a focal point of the GRIN lens can be less than 0.8 mm. The catheter can have a diameter of approximately 2 mm. The focal point can be configured to be in the target. The elongate catheter body can be an atherectomy catheter. The elongate catheter body can be an occlusion crossing catheter having a rotatable tip and a guidewire lumen therein. 
     In general, in one embodiment, a method of imaging a target includes inserting a catheter into a lumen of the target; transmitting optical radiation from a source through an optical fiber and a graded index fiber within the catheter; transmitting the optical radiation from the graded index fiber through an interface medium; transmitting the optical radiation reflected from the target and reflected from a reference interface along a common path in the graded index fiber and the optical fiber to a detector; and generating an imaging of the target based upon the reflected optical radiation. 
     This and other embodiments can include one or more of the following features. A distance from an edge of the catheter and a focal point of the graded index fiber can be less than 0.8 mm. The focal point can be in the target. 
     In general, in one embodiment, a method of making a graded index fiber for use with a common path optical coherence tomography system includes selecting a required distal angle of the graded index fiber such that a reference interface at the distal end of the graded index fiber will produce a reference reflection of between −28 and −42 dB; and polishing or cleaving the distal end of the graded index fiber to the selected angle within a tolerance of less than 0.2°. 
     This and other embodiments can include one or more of the following features. The method can further include splicing a graded index fiber to a distal end of a single mode optical fiber. The method can further include placing the distal end of the graded index fiber in optical contact with an interface medium to form the reference interface. The method can further include making an external surface of the interface medium concave. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG.  1    shows an example of a prior art OCT system. 
         FIG.  2 A  shows an exemplary OCT system as described herein. 
         FIG.  2 B  shows the travel of light through and out of a GRIN fiber. 
         FIG.  2 C  shows an exemplary graph of the diameter of a light beam as it exits a standard optical fiber vs. exiting a GRIN fiber. 
         FIG.  2 D  shows the resolution of an OCT image. 
         FIG.  3 A  shows an exemplary graph of noise in an OCT detector vs. power. 
         FIG.  3 B  shows an exemplary graph of a breakdown of the types of noise contributing to the total noise in the graph of  FIG.  3 A . 
         FIG.  3 C  shows a chart including data drawn from the graphs in  FIGS.  3 A and  3 B . 
         FIG.  3 D  shows the angles associated with a light beam at an interface of two separate mediums with different refractive indices. 
         FIG.  4 A  shows an exemplary optical fiber having a GRIN fiber cleaved/polished at an angle and the tip inserted into a medium of known refractive index. 
         FIG.  4 B  shows the light path in the system of  FIG.  4 A  as it is reflected from a reflective surface in a catheter. 
         FIG.  5 A  shows the use of an optical fiber with a GRIN lens in an OCT catheter. 
         FIG.  5 B  shows a close-up of the distal end of the GRIN fiber of  FIG.  5 A . 
         FIG.  6 A  is a top view of an exemplary mirror at the distal tip of an OCT catheter. 
         FIG.  6 B  is a cross-sectional side view the embodiment of  FIG.  6 A . 
         FIG.  7    shows a medical (intravascular) catheter system equipped with an OCT system. 
         FIG.  8    shows a system for implementing the OCT system and catheter. 
         FIG.  9    shows one example of an optical circuit. 
         FIG.  10    is a schematic of an OCT system as described herein. 
     
    
    
     DETAILED DESCRIPTION 
     The Optical Coherence Tomography (OCT) catheters and systems described herein are configured to provide image guided intra-vascular procedures that may be particularly useful for the diagnosis and/or treatment of arterial disease. The systems may include a catheter, an umbilical connection, and a console. The system uses OCT to form an image of the intravascular environment close to the catheter cutter. During intraluminal procedures, such as atherectomy, problems can arise as a result of failure to properly identify target tissue. By using a catheter having a common path optical fiber for OCT, proper identification of target tissue can be improved. 
     In general, the catheters described herein include a common-path system with a graded index (GRIN) fiber attached to the distal tip of a single mode optical fiber in the catheter so as to act as a lens for focusing light. The distal portion of the GRIN fiber is modified to provide a reference reflection between −28 dB and −45 dB, such as −28 dB and −42 dB, −32 dB and −45 dB, or −32 and −42 dB. In some embodiments, a secondary reflection of optical radiation from an interface between the optical fiber and the graded index fiber is configured to be less than −60 dB. An outer diameter of the graded index fiber and a protecting coating around the graded index fiber can be less than 0.01 inches. This common path OCT system with GRIN fiber advantageously achieves focusing of the light beam while still achieving the desired stable reference reflection and maintaining a small profile of the catheter. 
     Referring to  FIG.  2 A , a common-path OCT system  100  includes a laser source  102 , such as a swept frequency light source. An optical fiber  104 , such as a single mode optical fiber, transfers radiation from the laser source  102  towards the target  114 . A graded index fiber (“GRIN fiber”)  222  is attached to the distal end of the optical fiber  104 . Further, the GRIN fiber  222  is in optical contact with an interface medium  106 , such as an adhesive, epoxy, or cement. The light exiting the GRIN fiber  222  is thus in optical contact with the interface medium  106  such that the light exiting the GRIN fiber  222  and entering the interface medium  106  sees only a single interface. In some embodiments, the GRIN fiber  222  is embedded in the interface medium  106 . 
     In the common-path OCT system  100 , the index of refraction of the interface medium  106  is different than the index of refraction of the distal edge of the GRIN fiber  222 . Because the indices of refraction of the interface medium  106  and the distal edge of the GRIN fiber  222  are different, a Fresnel reference reflection can be created, which can be used to generate the resulting OCT image. That is, part of the light from the light source will exit the GRIN fiber  222  to travel to the target  114  while part of the light will be reflected back from the distal end of the GRIN fiber  222  to form the reference reflection for the OCT system. Some of the light beam that exits the GRIN fiber  222  will encounter the target  114  and be reflected or scattered by the target  114 , and some of this scattered light will, in turn, reenter the distal end of the GRIN fiber  222  and interfere with the reference reflection. The interference signal then travels back through the fiber  104 . A Faraday isolation device  112 , such as a Faraday Effect optical circulator, can be used to separate the paths of the outgoing light source signal (i.e., the light being sent to the target from the light source) and the interference signals returning from the distal end of the fiber. The separated interference signal can travel back to a detector  110  located at the proximal end of the optical fiber  104 . The interference signal detected by the detector  110  can then be used create an image of the target, as described below. 
     Because the reflected or scattered target light in the OCT system  100  travels a longer distance than the Fresnel reflected reference light, the reflected or scattered target light can be displaced by frequency, phase and or time with respect to the reference beam. For example, if swept-source radiation is used, then the light from the target will be displaced in frequency. The difference in displacement in phase, time or frequency between the reflected or scattered target light and the reference light can be used to derive the path length difference between the end of the optical fiber tip and the light reflecting or light scattering region of the target. In the case of swept source OCT, the displacement is encoded as a beat frequency heterodyned on the carrier reference beam. Embodiments such as the system described with respect to  FIG.  2 A , where the light paths in the reference and signal arms are common, are called common path interferometers. Using common path interferometry helps provide a low cost disposable catheter, as it eliminate the separate reference arm without adding significant size or system requirements to the catheter itself. 
     Referring to  FIG.  2 B , the GRIN fiber  222  can be spliced, such as fusion spliced, onto the distal end of the optical fiber  104 . As shown in  FIG.  2 B , the core  224  of the GRIN fiber  222  is much larger than the core  231  of the optical fiber  104  in order to allow the light to be expanded and contracted properly within the core of the GRIN fiber  222 . For example, the diameter of the core  224  of the GRIN fiber  222  can be greater than 40 μm, such as greater than 50 μm, such as greater than 60 μm, such as 40-110 μm, e.g., approximately 100 μm, while the diameter of the core  231  of the optical fiber  104  can be less than 20 μm, such as less than 10 μm, such as 8-10 μm, e.g., approximately 9 μm. The GRIN fiber  222  can be spliced in such a manner so as to minimize any return loss caused by the transition from the optical fiber  104  to the GRIN fiber  222 , i.e., to prevent a ghost image from forming due to a secondary reflection caused at the interface between the optical fiber  104  and the GRIN fiber  222 . For example, the optical fiber  104  and the GRIN fiber  222  can be spliced together at an angle. In one embodiment, the secondary reflection cause at the interface between the optical fiber  104  and the GRIN fiber  222  is less than or equal to −60 DB. 
     The GRIN fiber  222 , in contrast to standard lenses and even traditional GRIN lenses, can be specifically constructed such that the GRIN fiber  222  (including the core and any cladding, such as a cladding of less than 15 μm, e.g., approximately 12.5 μm) is approximately the same outer diameter as the optical fiber  104  (including the core  231  and cladding 233), thereby helping to maintain a low profile for the common path OCT system. For example, in one embodiment, the combined outer diameter of the optical fiber  104  and the GRIN fiber with a protective coating (as described further below) are both less than 0.01 inches, such as less than 0.08 inches, such as approximately 0.0065 inches with a protecting coating thickness of greater than or equal to 0.007 mm, such as 0.015 mm, 0.02 mm, or 0.022 mm. Advantageously, the GRIN fiber  222 , in contrast to a larger focusing lens, can advantageously be substituted in place of a distal portion of an optical fiber in a catheter without having to make significant changes to the catheter design due to the relatively small diameter of the fiber. 
     Referring to  FIGS.  2 B and  2 C , the GRIN fiber  222  can advantageously help focus light coming out of the optical fiber  104 . That is, as shown in  FIG.  2 C , light coming out of a standard optical fiber will spread out, thereby lowering the resolution of the resulting OCT image. In contrast, the light coming out of a GRIN fiber will stay at a much smaller diameter throughout the imaging depth of the target and thus increase the resolution of the resulting OCT image. That is, the azimuthal resolution (δ θ ) will be increased by using a GRIN fiber, which can help maintain the dynamic range with the depth. The azimuthal resolution is defined as minimal feature size which can be accurately represented by the imaging system, as shown in  FIG.  2 D . The azimuthal resolution is primarily dependent on the beam diameter of the investigating beam. The resolution is at least two times the diameter of the spot size. In one embodiment, the beam diameter of the GRIN fiber at focus (which can be less than 1 mm from the catheter tip, as described further below) is 35±5 microns in contrast to a standard single-mode optical fiber where the light expands outward without focusing (as shown in  FIG.  2 C ). As a result, the minimal resolvable feature size can be reduced to less than 100 microns, such as approximately 70 microns. 
     Referring back to  FIG.  2 B , the GRIN fiber  222  can be cleaved at the distal end such that the length of the GRIN fiber  222  allows the light  203  coming out of the GRIN fiber  222  to focus at the required distance or focal length within the area being imaged. For example, the light can focus between 0.5 mm and 1.5 mm, such as approximately 1.1 mm away from the distal tip of GRIN lens. The beam diameter at the focus can be somewhere between 10 and 60 microns, such as between 20 and 40 microns. As described below, the GRIN fiber  222  can be placed within a catheter such that the focal point of the fiber  222  is directly within the target being imaged. 
     In some embodiments, as shown in  FIG.  2 B , a coating  207 , such as a polyimide coating, covers the optical fiber  104  to strengthen the fiber, such as to withstand the tortuosity within the catheter, and/or protect the fiber  104  from environmental factors such as water vapor. During splicing of the optical fiber  104  and the GRIN fiber  222 , part of the coating  207  near the distal end, such as approximately 5 mm from the distal end, of the optical fiber  104  can be removed in order to allow for proper splicing, e.g., to allow the splicing equipment proper access to the fiber  104  and/or to prevent the coating  207  from melting over the distal end of the fiber  104  during splicing. In order to strengthen and protect the optical fiber  104  after splicing, the fiber  104  can be recoated with a recoat layer  209 , which can be, for example, polyimide. 
     Referring back to  FIG.  2 A , the laser source  102  can operate at a wavelength within the biological window where both hemoglobin and water do not strongly absorb the light, i.e. between 800 nm and 1.4 μm. For example, the laser source  102  can operate at a center wavelength of between about 1300 nm and 1400 nm, such as about 1310 nm to 1340 nm. The optical fiber  104  can be a single mode optical fiber for the ranges of wavelengths provided by the laser source  102 . Further, the gradient index profile of the GRIN fiber can be chosen such that the desired spot size is achieved at the desired focal distance. Furthermore, the distal end of the GRIN fiber  222  and the interface medium  106  can have specifically-chosen indexes of reflection such that a known magnitude of reference reflection is created. For example, the indexes of reflection can be chosen such that noise in the OCT system is minimized. 
     Noise in OCT systems comes from at least three sources: shot noise, thermal or Johnson noise, and residual intensity noise (RIN noise). There may additionally be noise from the analog-to-digital conversion process. RIN noise comes from noise intrinsic to the light source, tends to dominate at high reference powers, and can be limited by limiting the maximum laser light intensity, working with an alternative low RIN light source (non-laser), or by using balanced detection. Thermal (Johnson) noise tends to dominate at low reference power levels, and can be avoided by working at reference power levels yielding a DC photodiode current above that of the thermal noise floor. 
     Shot noise dominates in between RIN noise and thermal (Johnson) noise. Shot noise is caused by statistical fluctuations in the number of photons or electrons that carry a particular signal. For a well designed system, shot noise is the limiting factor in dynamic range. The indexes of refraction of the GRIN fiber  222  and the interface medium  106  can thus be chosen such that the OCT system  100  operates close to the shot noise limit. 
     The shot noise limit of a particular receiver is set by the responsivity of the photodetector, the detection bandwidth desired, and the reference DC power impinging on the detector element. An exemplary graph of a noise v. power is shown in  FIG.  3 A  with a break-down by the type of noise shown in  FIG.  3 B . The graphs in  FIGS.  3 A and  3 B  assume a system having 10 mW of forward power, 1550 nm center wavelength, 20 nm bandwidth, 1 MHz detection bandwidth, and a 1 A/W responsivity. 
     The shot noise limit is the area  301  at the bottom of the curve in  FIG.  3 A , at which the noise is the lowest or where the degradation from the shot noise limit is the least. Using the graph for a particular receiver, such as the graphs shown in  FIG.  3 A  and  FIG.  3 B , the desired power at the detector, P det , can be determined that would place the noise within a desired range of the shot noise limit. For example,  FIG.  3 C  shows a table of values drawn from  FIG.  3 B . Referring to  FIG.  3 C , a power of 0.158 μW would place the receiver at the minimum degradation point, 2.36 dB above the shot noise limit. Moreover, reference powers of between 63.1 nW and 251 nW would place the noise within 3 dB of the shot noise limit. Reference powers of between about 25 nW to 0.631 μW would place the noise within 5 dB of the shot noise limit. 
     To determine the total power, P out , that must be reflected from the interface  106  to obtain the desired P det , the losses of the detector  110  must be taken into account according to Equation 1:
 
 P   det   =P   out (1− L )   (equation 1)
 
where P out  is the power reflected from the reference interface and L is the sum of the optical losses from the distal end of the probe to the detector  110 . Therefore, assuming that P det  is equal to 0.2 μW (rounding from the 0.158 μW determined to place the noise as low to the shot noise limit as possible) and that the intermediate optical system operates at 90% efficiency such that L is 10%, then P out  is equal to 0.2 μW/(0.9)=0.2222 μW.
 
     The forward power at the distal end of the GRIN fiber prior to entering the interface medium is given by P in . In one exemplary embodiment, P in  can be equal to 10 mW. 
     Moreover, P out  and P in  can be used to determine the reflectivity of the reference interface  180 , according to equation 2:
 
 P   out   =P   in   R   2    (equation 2)
 
where R is the Fresnel coefficient of reflectivity. Therefore, assuming that P out  is 0.2222 μW and P in  is 10 mW (as described above), then R is equivalent to 0.004714.
 
     Moreover, the Fresnel equation (shown by equation 3) governs the intensity of reflection from a normal or near normal interface: 
                   R   =     (         n   1     -     n   2           n   1     +     n   2         )             (     equation   ⁢           ⁢   3     )               
where the index of refraction of the transparent medium is given by n 2  and that of the core is n 1 .
 
     If the distal end of the GRIN fiber  222  is kept polished such that it has a normal interface with the interface medium, then the refractive index of the GRIN lens, n 1 , will be fixed and can be assumed to be substantially equal to the refractive index at the center of the GRIN lens. For example, if the refractive index at the center of the GRIN lens is 1.4677, then the refractive index of the interface medium n 2  would have to be 1.4816 or 1.4539, according to equation 3 above. Thus, an interface medium of either index will produce the desired reference reflection. In some embodiments, the medium with the higher index of refraction may be preferable as it may be more readily available and/or have better mechanical properties, such as tensile strength. 
     The interface medium used with system  100  can be, for example, an adhesive. Depending upon the required index of refraction, the interface medium can be, for example, M21-CL which is a thermal curing adhesive. Another exemplary interface medium is the Light Weld® UV curable photonics adhesive OP-4-20658, produced by Dymax corporation, Torrington Conn. This adhesive, which has a refractive index of 1.585 in the cured state, is a rigid clear UV-curable adhesive that can be applied in a liquid form, and which then cures to a rigid form within seconds of exposure to UV light. Another exemplary transparent medium is EpoTek OG127-4 or OG116, produced by Epoxy Technology, Billerica Mass. This has a refractive index of 1.602 in the cured state. Another exemplary transparent medium is Masterbond EP42HT-2, which has a cured refractive index of 1.61. Another exemplary transparent medium is Norland Optical Adhesive NOA-61, which has a refractive index of 1.56 upon curing. 
     If an interface medium having the exact refractive index desired cannot be found (for example because it does not have the proper tensile strength or is not biocompatible), an interface medium having a refractive index that is close can be selected and the power in, P in , can be adjusted accordingly. Using the known r and the desired power at the detector, P det , the required power in P in  can then be determined according to equation 4:
 
 P   det   =P   in   R   2 (1− L )   (equation 4)
 
     In some implementations, the interface medium can be applied in a semi-liquid state, such as by dispenser, ink jet deposition, spraying, painting, dipping, or other process. The medium may then be cured to a solid form, such as by UV curing, thermal curing, chemical curing, drying, or other process. Other processes, such as vacuum deposition of transparent medium or direct mechanical placement of the transparent medium may also be used. 
     The interface medium can have a minimum thickness (i.e. depth between the end of the optical fiber and the end of the interface medium) of at least 
                 λ   min       2   ⁢   π       ,         
where λ min  is the wavelength of light in the optical fiber. For a wavelength of over 1250 nm, this will be approximately 200 nm or greater. The interface medium can also have a thickness that is great enough to introduce an offset between the reference reflection and the minimum distance that the target can approach the distal exit face of the GRIN fiber  222 .
 
     Alternatively, if a particular interface medium is desired due, for example, to the adhesive properties and/or biocompatibility of the interface medium, then the desired reference reflection can be achieved by cleaving or polishing the distal end of the GRIN fiber at a particular angle. In this case, equation 5 governs the reflection from the surface assuming that the incident light is unpolarized and therefore contains an equal mix of s and p polarizations: 
                   R   =              R   S     +     R   P       2                  (     equation   ⁢           ⁢   5     )               
where Rs is the reflection coefficient caused by the s-polarized light and is determined by equation 6:
 
                     R   S     =                n   1     ⁢   cos   ⁢           ⁢     θ   i       -       n   2     ⁢   cos   ⁢           ⁢     θ   t               n   1     ⁢   cos   ⁢           ⁢     θ   i       -       n   2     ⁢   cos   ⁢           ⁢     θ   t                        (     equation   ⁢           ⁢   6     )               
where is θ i  is the angle of incidence and θ t  is the angle of transmittance (see  FIG.  3 D ). Further, as shown in  FIG.  3 D , θ r  is the angle of reflection. Further, Rp in equation 5 is the reflection coefficient caused by the p-polarized light and is determined by equation 7:
 
     
       
         
           
             
               
                 
                   
                     R 
                     P 
                   
                   = 
                   
                      
                     
                       
                         
                           
                             n 
                             1 
                           
                           ⁢ 
                           cos 
                           ⁢ 
                           
                               
                           
                           ⁢ 
                           
                             θ 
                             t 
                           
                         
                         - 
                         
                           
                             n 
                             2 
                           
                           ⁢ 
                           cos 
                           ⁢ 
                           
                               
                           
                           ⁢ 
                           
                             θ 
                             i 
                           
                         
                       
                       
                         
                           
                             n 
                             1 
                           
                           ⁢ 
                           cos 
                           ⁢ 
                           
                               
                           
                           ⁢ 
                           
                             θ 
                             t 
                           
                         
                         - 
                         
                           
                             n 
                             2 
                           
                           ⁢ 
                           cos 
                           ⁢ 
                           
                               
                           
                           ⁢ 
                           
                             θ 
                             i 
                           
                         
                       
                     
                      
                   
                 
               
               
                 
                   ( 
                   
                     equation 
                     ⁢ 
                     
                         
                     
                     ⁢ 
                     7 
                   
                   ) 
                 
               
             
           
         
       
     
     Accordingly, if the refractive index of the medium, n 2 , is fixed because a particular medium is desired, then the Fresnel reflection can be adjusted by changing the cleave or polishing angle of the distal end of the GRIN fiber. In some embodiments, the cleave or polishing angle can be cleaved or polished to the desired angle within a tight tolerance, such as a tolerance (or variation from the desired angle) within less than 0.5°, less than 0.4°, less than 0.3°, or less than 0.2°, such as less than or equal to 0.15°. In contrast to standard single mode fibers, cleaving or polishing the distal end of the GRIN fiber within such tight tolerances is important to obtain the desired index of reflection because the back-reflection in GRIN fiber is much more sensitive to the cleave angle than the back-reflection in standard single-mode fiber. 
     Thus, as shown in  FIG.  4 A , the GRIN fiber  222  can be cleaved or polished at the distal end  401 , thereby providing the appropriate Fresnel coefficient of reflectivity. For example, the angle can be less than 4 degrees, less than 2 degrees, or less than 1 degree (plus or minus the tight tolerance as described above). In one embodiment, the combination of the angle between the interface medium  106  and the GRIN fiber  222  and the mismatch between the two refractive indices gives −28 dB and −45 dB, such as −28 dB and −42 dB, −32 dB and −45 dB, or −32 and −42 dB return loss. 
     Once the approximate angles have been determined using the equations described herein, the polishing can be adjusted slightly to compensate for any loss of light caused by the interface between the GRIN fiber and the optical fiber. That is, the Fresnel equations described herein do not take into account the interface between the GRIN fiber and the optical fiber. If the distal end of the GRIN fiber is polished or cleaved at an angle, then the coupling efficiency of resulting beam of reflected light traveling through the GRIN fiber and into the Single Mode optical fiber may change. Thus, while the distal angle required can be approximately determined by the equations described herein, further adjustment may be required in order to achieve the desired reflection. Further, the tolerance on the set angle for required reference reflection is very high as a the magnitude of change in reference reflection is effected both by Fresnel Reflection and coupling efficiency of the light from GRIN fiber into the Single Mode optical fiber. In some embodiments, additional factors, such as the diameter of the GRIN fiber, can also be adjusted to achieve the desired reflection. The device described herein can thus include a GRIN fiber as a focusing element that generates a stable reference reflection at the distal tip of the GRIN fiber. 
     Referring to  FIG.  4 B , a reflective surface  180 , such as an angled mirror (e.g., angled at 35-55 degrees relative to the axis of the optical fiber, such as 45 degrees), can be used to deflect the  203  beam in the desired orientation (such as into adjacent tissue). As shown in  FIG.  4 B , the focus of the beam  203  can be within the tissue (after reflecting off of the reflective surface  180 ). 
     Referring to  FIGS.  5 A and  5 B , the imaging system described herein can be used with a catheter, such as an atherectomy catheter or an occlusion crossing catheter. The atherectomy catheter can have a cutting tip and a tissue storage chamber. The occlusion crossing device can have a rotatable tip and a guidewire lumen. Exemplary catheters with which the imaging systems described herein can be used are described in copending Patent Applications: U.S. patent application Ser. No. 12/829,267, titled “CATHETER-BASED OFF-AXIS OPTICAL COHERENCE TOMOGRAPHY IMAGING SYSTEM”, filed Jul. 1, 2010, U.S. Pat. No. 9,125,562; U.S. patent application Ser. No. 13/433,049, titled “OCCLUSION-CROSSING DEVICES, IMAGING, AND ATHERECTOMY DEVICES”, filed Mar. 28, 2012, U.S. Pat. No. 8,644,913; 
     U.S. Provisional Patent Application No. 61/799,505, titled “OCCLUSION-CROSSING DEVICES”, filed Mar. 15, 2013; International Patent Application No. PCT/US2013/032679, titled “CHRONIC TOTAL OCCLUSION CROSSING DEVICES WITH IMAGING”, filed Mar. 15, 2013, Publication No. WO 2014/143064; U.S. patent application Ser. No. 12/829,277, titled “ATHERECTOMY CATHETER WITH LATERALLY-DISPLACEABLE TIP”, filed Jul. 1, 2010, U.S. Pat. No. 9,498,600; U.S. patent application Ser. No. 13/175,232, titled “ATHERECTOMY CATHETERS WITH LONGITUDINALLY DISPLACEABLE DRIVE SHAFTS”, filed Jul. 1, 2011, U.S. Pat. No. 9,345,510; U.S. patent application Ser. No. 13/654,357, titled “ATHERECTOMY CATHETERS AND NON-CONTACT ACTUATION MECHANISM FOR CATHETERS”, filed Oct. 17, 2012, Publication No. US-2013-0096589-A1; U.S. patent application Ser. No. 13/675,867, titled “OCCLUSION-CROSSING DEVICES, ATHERECTOMY DEVICES, AND IMAGING”, filed Nov. 13, 2012, U.S. Pat. No. 9,345,406; International Patent Application No. PCT/US2013/031901, titled “ATHERECTOMY CATHETERS WITH IMAGING”, filed Mar. 15, 2013, Publication No. WO 2013/172970; and International Patent Application No. PCT/US2013/032494, titled “BALLOON ATHERECTOMY CATHETERS WITH IMAGING”, filed Mar. 15, 2013, Publication No. WO 2014/039099, all of which are incorporated by reference in their entireties. 
     As shown in  FIG.  5 A , the distal end of the optical fiber  104  and the GRIN fiber  222  can sit away from the central axis  555  of an imaging catheter  502  such that the distal tip of the GRIN fiber  222  is near the outer edge  525  of the catheter  502  (i.e., where the diameter of the catheter is the largest). Further, in some embodiments, the catheter  502  can be configured to be placed in a blood vessel of 1-8 mm in diameter, e.g., 2-5 mm, such as approximately 3 mm in diameter. The catheter  502  can thus have a diameter of approximately 2 mm. By placing the GRIN fiber  222  off-axis and near the outer edge of the catheter  502 , the focus can be located less than 1 mm from the outer edge  525  of the catheter  525 , such as less than 0.8 mm or less than or equal to 0.7 mm from the outer edge  525  of the catheter  502 , while still being located within the tissue of interest (e.g., the wall of the vessel). Thus, even after the light beam  203  bounces off of a reflective surface  180 , such as a mirror or a prism, the maximum portion of the focal range lies inside the vessel structure, as indicated by the arrows  517 . Having the maximum portion of the focal range lie inside the vessel structures improves both resolution and quality of the resulting image of the tissue. 
     Referring to  FIGS.  6 A and  6 B , the reflective surface  180  can be designed and optimized in a catheter  502  in order to fit into the small (approximately 2 mm) diameter of the catheter head and to reflect into a blood vessel tissue located up to 1-4 mm away from the side of the distal catheter tip. As shown in  FIG.  6 B , the catheter  502  can include a cut-out  2160  configured to hold the edge of the GRIN fiber  222  and the reflective surface  180 , such as a mirror or prism. In one embodiment, the reflective surface  180  can include a silicon die  401  having a reflective coating  403 . The reflective coating  403  can be, for example, a gold coating. The reflective coating  403  can be greater than 
               λ   min       2   ⁢   π           
thick, where λ min  is the wavelength of light in the optical fiber. For example, the metallic coating can be greater than about 2,800 Å thick.
 
     Further, the surface of the silicon die  401  under the reflective coating  403  can be polished to less than 400 nm peak-to-peak roughness, such as better than 300 nm peak-to-peak roughness, for example about 200 nm peak-to-peak roughness. An adhesive, such as nickel, titanium, or chromium, can be used to adhere the gold coating to the silicon die. The adhesive can be between about 50 Å and 200 Å thick, such as about 100 Å thick. The reflective surface  180  of this configuration can be at least 95% reflective, such as 98% reflective. 
     The reflective surface  180  can be placed on a slope such that it is at an angle of between 30° and 60°, such as 45° with respect to a longitudinal axis  405  of the core of the optical fiber  104 . Moreover, the reflective surface  180  can be configured such that the total distance that the light travels from the GRIN fiber  222  to the reflective surface  180  and out to the sample is between 100 and 400 μm, such as between 200 and 250 μm. 
     As shown in  FIGS.  6 A and  6 B , an opening  2610  can be formed in the catheter  502 , exposing the distal end of the GRIN fiber  222 . The OCT reflective surface  180  can be placed in the opening near the distal tip of the catheter  104 , and the interface medium can cover or embed the fiber  502  and opening  2610 . Further, referring to  FIG.  5 B , the interface medium  106  can have an outer surface  539  (i.e. along the outer edge  525  of the catheter  502  closest to the target) that is concave, i.e., that forms a meniscus within the cut-out  2160 . By having an outer surface  539  that is concave, the light beam  203  will hit the surface of the interface medium  106  at an angle (shown by the tangential line  215 ) such that any light that reflects back from the surface  539  will advantageous reflect at angles that are away from the distal end of the GRIN fiber  222 . Having the majority of the reflected light stay away from the distal end of the GRIN fiber  222  advantageously ensures that the light does not couple back into the fiber  222 / 104  to form secondary reflection and subsequently create mirroring images on top of each other with an offset (sometimes termed “ghost images”). Because the GRIN fiber has a core diameter that is much larger than the core diameter of a traditional single mode optical fiber, the GRIN fiber has higher acceptance angle and thus is more prone to create secondary reflection. Accordingly, the concave surface  525  is especially advantageous when using a GRIN fiber  222 . 
       FIG.  7    shows an overview of the main components of an OCT imaging system  500  including a fiber optic catheter  502 . The catheter  502  can be sized to fit into a blood vessel, e.g. can be about 2 mm in diameter. In this configuration, the OCT optical apparatus  504  (including the light source, optical circulator, and detectors) can be located at the proximal end of the catheter  502 , and can be connected to an image processor and a display  506 . The distal end of the catheter  502  includes the image fiber and the mirror. The system  500  is designed to be used within the body of a patient for various medical purposes, such as occlusion crossing or atherectomy. Thus, other components, such as a vacuum  510 , aspiration control  508 , and a debris reservoir  512  may be useful. 
     The system described herein may be used to produce relatively narrow angle images of a portion of an interior lumen of a human body, such as the interior of a blood vessel. Looking at a section of a tissue through a single OCT optical fiber is limited in that the useful angle of view produced by a single OCT optical fiber is at most a few degrees. In order to produce a more medically useful panoramic view of a wide arc or swath from the interior of a blood vessel, such as 45°, 90°, 120°, or more, the catheter containing the optical fiber can be rotated. 
     The system described herein can produce images, e.g. images of tissue morphology, having an axial resolution of around 6-15 microns, e.g. 8-10 microns, and to depths of 1-2 mm depending on the optical properties of the sample being imaged. The axial resolution of the OCT system can be about ten times higher than that of a similar ultrasound system. The azimuthal resolution can be maintained to be less than 100 microns typically up to depths of 3-4 mm. 
       FIG.  8    shows a system  2700  for implementing the OCT system and catheter described herein. A power supply  2713  supplies power to the OCT engine  2703 , the computer processor  2707 , and the optical system  2711 . A trigger  2701  in the OCT engine  2703  is connected to a trigger  2705  in the computer processor  2707  to begin processing of the image. Moreover, a catheter rotation encoder  2715  is attached to the computer processor  2707  to transfer signals related to the location and rotation of the optic fiber. The OCT detector  2717  is attached to the computer processor  2707  to process the final image. Finally, a video signal is sent from the computer processor  2707  to a monitor  2709  to output the image to the user. 
     In some embodiments, the OCT system and catheter described herein can image up to 1-2 mm in depth with axial resolutions around 8-10 microns, sufficient to give the physician highly detailed images almost to the cellular organization level and visibility beyond the maximum cut range of the catheter. Moreover, the OCT atherectomy catheter described in can advantageously have imaging capability with crossing-profile impact that is much smaller than traditional OCT systems and ultrasound transducers. 
     In one example, an image-guided interventional catheter (e.g., an OCT catheter as described above) may be used to address unmet needs in peripheral and coronary artery disease (atherosclerosis). The system may include a console having a modest footprint and in a cath lab without need for extensive integration into cath lab systems. In some variations, the systems described herein may be integrated with other catheter (e.g., guidance, control, imaging) systems. The system may be configured to allow a procedure to start/proceed/finish under fluoro guidance in the event of a system failure. The system is also configured to be compatible with sterile procedures. 
     As mentioned above, the OCT systems described herein may allow real-time information on intravascular lesion morphology and device orientation in the vessel. This and other features may also allow improved navigation precision around complex anatomy (e.g., bifurcations, ostials, tortuosity, cutting on a curve, etc.), and around stent struts. The catheters may be safely used to traverse diseased tissue while reducing incidence of perforations and dissections potentially associated with a more aggressive treatment strategy. The systems may also provide immediate assessment of acute procedural success, and a reduction in procedure time compared to contemporary interventional techniques. The systems described herein may allow imaging of vessel wall morphology in real time and at a level of precision that could assist the physician in making a “diseased/not-diseased” determination. 
     In one example, the OCT system is configured to allow tissue morphology to be imaged in real time with resolution routinely around 8-10 microns, and to depths of 1-2 mm depending on the optical properties of the tissue. The axial resolution of OCT is sufficiently high that the images presented to the operator substantially resemble histology from optical microscopy, and are as a result more intuitively interpreted than ultrasound or MRI/CT images. The depth to which OCT can image through tissue with minimal to moderate lipid content is sufficient to give the physician visibility beyond the maximum proposed depth of cut for an atherectomy catheter, allowing the safety margins of the putative cut to be assessed. 
     As mentioned, OCT has several other technical and economic advantages for catheter applications. The impact on catheter crossing profile of the OCT optical fiber is much smaller than for even the smallest comparable ultrasound transducer. The axial resolution of OCT is typically 10× higher than ultrasound; this translates directly to image interpretability. The limited depth of penetration of typical OCT devices is not of primary concern in this application in many applications, because it is known from prior atherectomy procedures that substantial clinical benefit can be obtained by removing several hundred micron thicknesses of tissue. The depth of penetration may be matched to the expected maximum cut depth. Regions of particularly deep or thick tissue (target tissue to be removed) may be identified and treated serially or separately. For example, highly lipid-rich tissues (necrotic cores) appear as dark voids in OCT images, typically with bright caps. 
     The center wavelength for the optical system may be chosen to provide sufficient depth of penetration, as well as compatibility with the components of the system. For example, the OCT systems may use light that can be transmitted through fused silica fiber optics (where the primary investment in cost and quality has been made). The wavelength range to 250-2000 nm may be particularly useful. Single mode fibers can be readily obtained at any of these wavelength ranges, although wavelengths above 400 nm may be preferable. Below 250 nm air-guiding fibers can also be used. It may be easier to “see” through small annuli of either blood, saline or mixtures by restricting the scan range of the source to regions where hemoglobin and water do not strongly absorb light. This leads to the use of a “biological window” between about 800 nm and 1.4 microns. 
     The dominant mechanism restricting penetration depth in biological tissue when using ballistic optical scattering techniques is the photon scattering cross-section in the tissue. Higher scattering cross-sections causes fewer photons to traverse from source to target and back ballistically, that is with only one scattering event at the target leading to a reduction in useful signal. The scattering cross-section scales as an inverse power of wavelength over the 250-2000 nm range, transitioning from an exponent of −4 at shorter wavelengths to a smaller value at longer wavelengths. The value decreases monotonically going from short to longer wavelengths. Therefore, if it is desired to see deeper in tissue, the wavelength range of the source should be biased to longer wavelengths. Moving to longer wavelengths may, in some embodiments, require a more sophisticated laser source to achieve the same resolution compared to imaging at shorter wavelengths, however this is a soluble technical problem. 
     In some variations, the system takes advantage of the widespread availability of cheap, high quality parts. For example, fiber-based telecommunications has evolved at three specific center wavelength ranges; 800 (LAN only), 1310 (O-band) and 1550 nm (C-band). The systems described herein may use a center wavelength to 1310 nm, though this does not mean that the other two wavelength ranges could not be made to work. For example, the 800 nm center wavelength range is routinely used in ophthalmology, where depth of penetration can be sacrificed for tissue layer resolution and where fiber delivery is not a requirement (free-space optics may be used). In some variations, the system works in the telecommunications O-band. In practice the range of center wavelength is 1315-1340 nm may be dictated by the availability of suitable laser sources in the O-band. 
     There are three primary categories of source/detector combinations in OCT, namely Time-Domain, Spectral-Domain (Fourier Domain or Spectral Radar) and Swept Source OCT. The examples of OCT systems described herein are swept source OCT (SS-OCT), which allow for video-rate imaging, few or no moving parts, a simple optical system suitable for fiber implementation, imaging to depths greater than 1 mm, and insensitivity to the rigors of a mobile environment. 
     As discussed above, several interferometer configurations may be used. The systems described herein are Common Path Interferometry (CPI) systems. This has several advantages given the goal of catheter based imaging with cost-constrained capital equipment and disposable devices. The SS-OCT with CPI system described herein preserves the Fellgett Advantage. Fellgett&#39;s advantage or the multiplex advantage is an improvement in spectroscopic techniques that is gained when an interferometer is used instead of a monochromator or scanning delay line. The improvement arises because when an interferometer is employed, the radiation that would otherwise be partially or wholly rejected by the monochromator or scanning delay line in its path retains its original intensity. This results in greater efficiency. This embodiment contrasts with the other systems, in which only a small fraction of the laser power is useful at any given time. For example, the Lightlab™ M2 system uses TD-OCT with a scanning delay line, which is equivalent for the purposes of the Fellgett Advantage to a monochromator. Clinically, the Fellgett advantage impacts imaging speed (frame update rate), allowing significant improvements in video display rates which translate to a reduction in ambiguity in interpreting the image. 
     The CPI systems described herein also preserve the Jacquinot Advantage. The Jacquinot advantage states that in a lossless optical system, the brightness of the object equals the brightness of the image. Assuming that losses due the optical components are negligible, an interferometer&#39;s output will be nearly equal in intensity to the input intensity, thus making it easier to detect the signal. This translates directly to image quality, and a more interpretable image. 
     The CPI system as described herein therefore makes highly efficient use of the laser power. Light is either used for the reference reflection or impinges on the tissue and is used to create signal. No light is lost in attenuators or additional optical components or unused reciprocal paths. This efficient use of laser power is most apparent in the ability of the system to display clinically relevant images of the intravascular environment in real time, without the need for extensive post processing or even on-the-fly image correction. 
     Furthermore, these systems are “down-lead insensitive”, allowing the connection from catheter to console to be of almost arbitrary length without requiring a matched reference delay line to be shipped with each catheter. This minimizes the additional cost impact of the imaging components added to the catheter. It also allows a console component to be positioned almost anywhere, minimizing the potential disruption to work flow and minimizing the threat to a sterile field. 
     The systems described herein also minimize the number of optical components in the imaging system which could contribute to chromatic aberration. This minimization preserves the spectral fidelity of the laser source optimizing the layer resolution. This translates directly to image quality, and a more interpretable image. 
     The common-path systems described herein also have exceptional phase stability. Path length changes affecting the sample arm (temperature changes, stress-induced birefringence etc) also affect the reference arm identically. The distance from the ZPD (zero-pathlength difference) point (the reference plane) to the sample is physically fixed and is not subject to variability due to turbulence. This exceptional phase stability coupled with the exceptional phase stability of the OCT engine means that the Z-axis of the display (depth) has minimal jitter, in turn maximizing the real-time interpretability of the image. It also allows mathematical manipulation of the data that would otherwise be impossible. For example, one advantage of the systems described herein is the ability to perform pre-FFT averaging, which lowers the overall noise floor of the system again translating directly to image quality and interpretability. 
     In one example, the catheter is around 2 mm in diameter (7F compatible). In a saline-filled lumen, the system will be able to detect an interface (e.g., vessel wall) at 2 mm from the OD of the catheter. In this variation, the following parameters may be used for the catheter and system: 
     
       
         
           
               
               
               
             
               
                   
                   
               
               
                   
                 Specifications 
                 Value 
               
               
                   
                   
               
             
            
               
                   
               
            
           
           
               
               
               
               
            
               
                   
                 Optimized Detector Bandwidth 
                 DC-10 
                 MHz 
               
               
                   
                 Nyquist/Shannon rate 
                 20 
                 MHz 
               
            
           
           
               
               
               
            
               
                   
                 Minimum number of points to 
                 630 
               
               
                   
                 sample for full resolution 
               
               
                   
                   
               
            
           
         
       
     
     The detector may detect optical modulation on the carrier wave from DC to at least 10 MHz with no roll-off in sensitivity. To prevent aliasing (which complicates image interpretation) we may digitize the detector output at a minimum of 20 M-Samples/sec (Nyquist limit) to preserve interpretable real time imaging capability. We may thus capture at least 630 points per laser pulse at this digitizer rate to avoid undersampling the available laser bandwidth. 
     A practical resolution target is the intima of healthy coronary artery. The system resolution is capable of showing the intima (endothelial layer+internal elastic lamina) as a single sharp bright line on the display. 
     The system may have an impulse response of 8-10 microns. This resolution dictates the laser scan range requirements and the bandwidth requirements of all the optical components in the fiber harness through the equation: 
     
       
         
           
             
               ∂ 
               z 
             
             = 
             
               
                 
                   2 
                   ⁢ 
                   ln 
                   ⁢ 
                   
                       
                   
                   ⁢ 
                   2 
                 
                 π 
               
               ⁢ 
               
                 
                   λ 
                   0 
                   2 
                 
                 
                   n 
                   ⁢ 
                   Δ 
                   ⁢ 
                   λ 
                 
               
             
           
         
       
     
     Where δz is the axial resolution, λ is the wavelength, Δλ is the wavelength range over which the laser scans, n is the refractive index of the medium and the other symbols have their usual meaning. The origin of this relationship is the Heisenberg Uncertainty Principle. Several observations accrue from this equation. 
     If the laser scan range Δλ is not broad enough, δz (the resolution) is compromised and an image of a step refractive index discontinuity will be blurred out over many pixels. If any of the optical components in the system restrict (alternatively called clipping or vignetting) the effective bandwidth of the system is reduced and the resolution may suffer. Since the resolution equation has the center wavelength squared in the numerator, as we move to longer center wavelengths for the reasons described above, commensurately larger laser scan range may achieve equivalent axial resolution. Ophthalmology is routinely performed at 800 or 1000 nm center wavelength where there is no need to image deeply into the retina, but where the available lasers allow extremely high resolution of the layers of the retina (down to 1-2 microns thickness). 
     In some variations, the OCT system has a scan range of &gt;100 nm. The theoretical resolution of this engine is 6.35 microns in a medium with a refractive index of 1.35. Stipulating that we digitize at least at the Nyquist limit, fully sample the scanned bandwidth, and that the rescaling procedure in the software does not distort the data, the theoretical resolution of this system is sufficient to show the intima of a healthy coronary artery at the impulse response limit. 
     The choice of 1310 nm as a center wavelength for the laser means that we may use standard commercial off-the-shelf telecommunications components which have guaranteed performance at this wavelength and for which standardized test protocols exist. Reasonable and customary incoming inspection procedures can be used to verify that the components going into the system will not deteriorate image quality. 
     As mentioned above, the system may include receiving electronics including a detector. Assuming that the operating center wavelength is 1315-1340 nm with a full-width half maximum responsivity of &gt;100 nm, and that the detector operates as close as reasonably possible to the shot-noise limited regime, the system may have sufficient trans-impedance gain from the detector to allow the A/D card to operate at an input range where digitizer noise is not a dominant contributor to the noise floor of the system. 
     Manufacturing tolerances on the catheters will yield a range of distal tip reference reflection intensities. The detector may be configured or chosen so as not to saturate at the high manufacturing limit of the reference reflection power. In one example, the system uses a Fermionics FD80 photodiode in an FC receptacle package as the active element in the photodetector. 
     The system may also include a fiber harness designed to: 1) provide a low loss pathway from the laser to the catheter, 2) route signal light returning from the catheter to the detector, 3) allow the bleed-in of a red laser diode signal to allow rapid assessment of the integrity of the fiber from cable to distal tip, and 4) provide manufacturing, calibration and field service housekeeping signals to facilitate console production, validation and maintenance. 
     One primary component of the fiber harness may be a self-contained enclosure with bulkhead FC/APC receptacles on it and containing an optical circuit (such as the one shown in  FIG.  9   . 
     In one example, the fiber harness may be connected as: #1 Incoming OCT source (e.g., Santec) Santec output connected here. #2 Diagnostic port (OSA/Photodiode/MZI Calibration); #3 Diagnostic port (OSA/Photodiode/MZI Calibration); #4 Connection to Detector; #5 Reflected FBG Marker (Time/Wavelength Calibration Point); #6 Connection to Catheter; #7 Transmitted FBG Signal (Photodiode scope trigger); #8 Connection to red laser source. Connections may be made with single mode fiber with a cut-off of &lt;1260 nm. The inputs/outputs do not need to be optically isolated. 
     In some variations, an electrical harness may be used. The electrical harness may be configured to: 1) provide isolation for the various electrical components in the imaging system; 2) distribute 110V to the OCT engine, slave monitor and computer; 3) provide regulated isolated housekeeping power at appropriate voltages and amperages to the detector, red diode laser, catheter handle azimuthal position encoder; 4) send the video signal to the remote monitor; and 5) receive the catheter handle azimuthal angle encoder signal back to the console. 
     Line power may enter the console through a standard IEC60320 type C14 male power cord entry connector. The power cord used may be Hospital Grade and may have a standard IEC60320 type C13 female connector at the console end. An isolation transformer can distribute LINE power to the OCT engine, slave monitor and computer through IEC standard power cords. 
       FIG.  10    shows one example of a schematic of an OCT system as described herein. In this example, items with dotted perimeters are outside the main console chassis enclosure. Analog signal interconnects are to be made with RG58 (U, A/U) patch cables terminated with BNC connectors. The (Santec) Trigger Out signal is a falling edge signal (high Z) and should not be terminated in 50 ohms. The Encoder Signal can be terminated with a MiniCircuits low pass filter module at the A/D card to remove high frequency spurious noise. The Detector Signal can be terminated with a MiniCircuits low pass filter module at the A/D card to remove any noise in an irrelevant frequency range. 
     The optical fiber may have a cut-off less than 1260 nm and have single mode performance between 1270 and 1380 nm (and be manufactured compatible with SMF-28 standards). The mechanical connections (pigtail and patch cable) may include a simplex cable, and an inner loose tube Teflon Aramid fiber inner annulus to prevent stretching. The outer Jacket may be 2 mm polyurethane. The connector may be a Diamond E2108.6 connector with a 0.25 dB maximum insertion loss and a −65 dB maximum return loss. 
     The distal tip reference reflection (mirror) may include at least one (1) reflective interface, and may have a return loss of −33.5 dB (Nominal (31-35 dB)). There may be 200-250 microns solid transparent offset from interface to minimum tissue approach point. Interceding optical discontinuities between console and catheter distal tip may be kept to less than 65 dB return loss maximum for any individual surface. The number of reflective interfaces separated by less than 8 mm may be minimized. The parameters above are exemplary only, and may be varied as understood by those of skill in the art, while still remaining in the spirit of the invention as described herein. 
     The examples and illustrations included herein show, by way of illustration and not of limitation, specific embodiments in which the subject matter may be practiced. Other embodiments may be utilized and derived there from, such that structural and logical substitutions and changes may be made without departing from the scope of this disclosure. Such embodiments of the inventive subject matter may be referred to herein individually or collectively by the term “invention” merely for convenience and without intending to voluntarily limit the scope of this application to any single invention or inventive concept, if more than one is in fact disclosed. Thus, although specific embodiments have been illustrated and described herein, any arrangement calculated to achieve the same purpose may be substituted for the specific embodiments shown. This disclosure is intended to cover any and all adaptations or variations of various embodiments. Combinations of the above embodiments, and other embodiments not specifically described herein, will be apparent to those of skill in the art upon reviewing the above description. 
     When a feature or element is herein referred to as being “on” another feature or element, it can be directly on the other feature or element or intervening features and/or elements may also be present. In contrast, when a feature or element is referred to as being “directly on” another feature or element, there are no intervening features or elements present. It will also be understood that, when a feature or element is referred to as being “connected”, “attached” or “coupled” to another feature or element, it can be directly connected, attached or coupled to the other feature or element or intervening features or elements may be present. In contrast, when a feature or element is referred to as being “directly connected”, “directly attached” or “directly coupled” to another feature or element, there are no intervening features or elements present. Although described or shown with respect to one embodiment, the features and elements so described or shown can apply to other embodiments. It will also be appreciated by those of skill in the art that references to a structure or feature that is disposed “adjacent” another feature may have portions that overlap or underlie the adjacent feature. 
     Terminology used herein is for the purpose of describing particular embodiments only and is not intended to be limiting of the invention. For example, as used herein, the singular forms “a”, “an” and “the” are intended to include the plural forms as well, unless the context clearly indicates otherwise. It will be further understood that the terms “comprises” and/or “comprising”, when used in this specification, specify the presence of stated features, steps, operations, elements, and/or components, but do not preclude the presence or addition of one or more other features, steps, operations, elements, components, and/or groups thereof. As used herein, the term “and/or” includes any and all combinations of one or more of the associated listed items and may be abbreviated as “/”. 
     Spatially relative terms, such as “under”, “below”, “lower”, “over”, “upper” and the like, may be used herein for ease of description to describe one element or feature&#39;s relationship to another element(s) or feature(s) as illustrated in the figures. It will be understood that the spatially relative terms are intended to encompass different orientations of the device in use or operation in addition to the orientation depicted in the figures. For example, if a device in the figures is inverted, elements described as “under” or “beneath” other elements or features would then be oriented “over” the other elements or features. Thus, the exemplary term “under” can encompass both an orientation of over and under. The device may be otherwise oriented (rotated 90 degrees or at other orientations) and the spatially relative descriptors used herein interpreted accordingly. Similarly, the terms “upwardly”, “downwardly”, “vertical”, “horizontal” and the like are used herein for the purpose of explanation only unless specifically indicated otherwise. 
     Although the terms “first” and “second” may be used herein to describe various features/elements, these features/elements should not be limited by these terms, unless the context indicates otherwise. These terms may be used to distinguish one feature/element from another feature/element. Thus, a first feature/element discussed below could be termed a second feature/element, and similarly, a second feature/element discussed below could be termed a first feature/element without departing from the teachings of the present invention. 
     As used herein in the specification and claims, including as used in the examples and unless otherwise expressly specified, all numbers may be read as if prefaced by the word “about” or “approximately”, even if the term does not expressly appear. The phrase “about” or “approximately” may be used when describing magnitude and/or position to indicate that the value and/or position described is within a reasonable expected range of values and/or positions. For example, a numeric value may have a value that is +/−0.1% of the stated value (or range of values), +/−1% of the stated value (or range of values), +/−2% of the stated value (or range of values), +/−5% of the stated value (or range of values), +/−10% of the stated value (or range of values), etc. Any numerical range recited herein is intended to include all sub-ranges subsumed therein.