Abstract:
A system monitors or images a portion of a sample. The system includes an optical interferometer with a measurement arm, a reference arm, and an optical splitter. The arms are coupled to receive light from the optical splitter. One of the arms includes an acousto-optical modulator. The interferometer is configured to interfere light output from the two arms. The system also includes a detector that receives the interfered light and uses the received light to determine a depth-dependent quantity characterizing a portion of the interior of the sample.

Description:
[0001]     This application claims the benefit of U.S. Provisional Application No. 60/269,586, filed Feb. 17, 2001.  
       BACKGROUND OF THE INVENTION  
       [0002]     1. Field of the Invention  
         [0003]     This invention relates to optical monitoring and imaging.  
         [0004]     2. Discussion of the Related Art  
         [0005]     Contemporary medical technology uses x-rays, sound waves, and visible light to produce in vivo images of biological tissues. Visible light and infrared (IR) light imaging has a better potential resolution than imaging produced by sound waves, because visible and IR light has a shorter wavelength than sound waves. In spite of this advantage of visible and IR light, in vivo imaging systems often use sound waves, because visible and IR light does not penetrate thick tissues. Consequently, many in vivo imaging systems do not have the image resolutions obtainable in imaging systems based on visible or IR light.  
         [0006]     The resolution of in vivo examination systems is also limited by tissue motion. For many organisms internal tissue motions are always present, and these tissue motions interfere with tissue examinations that require more observation time than the time scale associated with the internal tissue motions. The internal tissue motions can cause image scans to produce multiple or smeared images. The internal tissue motions also cause the displacement of probes placed in the tissues. These displacements cause the probes to measure smeared out tissue properties.  
       BRIEF SUMMARY OF THE INVENTION  
       [0007]     Various embodiments include systems that process optical imagining data to obtain information on the position and/or velocity of a region inside a sample being imaged. In some embodiments, the position and/or velocity data is used to compensate for tissue motions during in vivo examinations of living organisms. The compensation corrects for smearing of scan images or measured properties that would otherwise result due to relative motions between the tissue and a monitoring probe.  
         [0008]     One system according to principles of the invention images a portion of a sample. The system includes an optical interferometer with a measurement arm, a reference arm, and an optical splitter. The arms are coupled to receive light from the optical splitter. One of the arms includes an acousto-optical modulator. The interferometer is configured to interfere light output from the two arms. The system also includes a detector that receives the interfered light and uses the received light to determine a depth-dependent quantity characterizing a portion of the interior of the sample.  
         [0009]     In some such systems, the detector uses the interfered light to determine a signed displacement or velocity of a portion of the sample. Such a system monitors body tissue movements and may further include a controller that receives displacement or velocity information. The controller adjusts data collection from the tissue in a manner that is responsive to changes in the tissue&#39;s location, orientation, or velocity relative to a monitoring probe. 
     
    
     BRIEF DESCRIPTION OF THE FIGURES  
       [0010]      FIG. 1A  shows a system that optically monitors or images a sample;  
         [0011]      FIG. 1B  is a flow chart for a process that uses the system of  FIG. 1A ;  
         [0012]      FIG. 2  shows a graded index (GRIN) fiber-size lens used in some embodiments of the probe of  FIG. 1A ;  
         [0013]      FIG. 3A  shows a conventional GRIN fiber lens;  
         [0014]      FIG. 3B  shows a refractive index profile for one embodiment of the GRIN fiber-size lens of  FIG. 2 ;  
         [0015]      FIGS. 4A and 4B  show embodiments of the system of  FIG. 1A  that incorporate optical interferometers;  
         [0016]      FIG. 4C  illustrates an acousto-optical modulator used in the interferometers of  FIGS. 4A and 4B ; and  
         [0017]      FIG. 5  shows a medical diagnostic system based on the system of  FIGS. 4A and 4C . 
     
    
     DETAILED DESCRIPTION OF THE EMBODIMENTS  
       [0000]     1. Optical Micro-Probe and Imagining System  
         [0018]     A co-pending patent application describes optical micro-probes and systems used in some embodiments of the invention of the present application.  
         [0019]      FIG. 1A  shows a system  10  for optically monitoring or imaging a region of a sample  12 , e.g., for endoscopic viewing of a biological tissue. Various embodiments of the system  10  determine the velocity and/or three-dimensional position of the region being monitored or imaged, e.g., via tomography. Such monitoring or imaging functions are useful for medical diagnostics and treatment, e.g., invasive imaging of anomalous tissue structures in vivo and monitoring of tissue motion during other medical procedures.  
         [0020]     The system  10  includes a source  14  of IR, visible, or ultraviolet light, an optical splitter or circulator  16 , an optical micro-probe  18 , and a light detector  20 . Exemplary sources  14  include monochromatic sources or multi-chromatic sources, e.g., a pulsed Ti-sapphire laser with a low coherence time of about 10 −15 -10 −13  seconds. The optical splitter or circulator  16  directs a portion of the light from the source  14  to the optical micro-probe  18 . The optical micro-probe  18  has a distal end  22  located either above or below the surface  23  of the remote sample  12 . The optical micro-probe  18  delivers the source light to a region of the sample  12 . The optical micro-probe  18  also returns to the splitter or circulator  16  a portion of the light scattered or emitted by the region of the sample  12  illuminated by the optical micro-probe  18 . The optical splitter or circulator  16  redirects the returned light to detector  20 . The detector  20  uses the returned light to determine a scattering or emission characteristic of the region of the sample  12  that produced the light. Some detectors  20  are configured to determine the distance of the region from the optical micro-probe  18  and/or the velocity of the region.  
         [0021]      FIG. 2  shows one embodiment  18 ′ of optical micro-probe  18  shown in  FIG. 1A . The optical micro-probe  18 ′ includes a single-mode optical fiber  24  that transports light to and from the sample  12 . The distal end  22  of the fiber  24  is fused to a GRIN fiber-size lens  26 , which has the same outer diameter as the optical fiber  24 . In some embodiments, the GRIN fiber-size lens  26  also has a rounded end face  28  that facilitates insertion of the end  22  of the optical micro-probe  18 ′ into samples such as biological tissues. In some embodiments, a portion of the GRIN fiber-size lens  26  adjacent the end face  28  has a conical taper (not shown). The taper also facilitates insertion of the optical micro-probe  18  into sample  12 , i.e., the taper functions like a needle point.  
         [0022]     The GRIN fiber-size lens  26  collimates light from fiber  24  into a focused beam  30 . The collimated beam  30  illuminates a region of the sample  12  located forward of the lens  26 . Points  32  in the illuminated region scatter or emit light in response to being illuminated. The backscattered or emitted light is useable for imaging or monitoring. The beam collimation enables resolving transverse locations of the points  32  with respect to the axis of the GRIN fiber-sized lens  26 , because points  32  producing backscattered or emitted light are located within the region illuminated by the beam  30 .  
         [0023]     In some embodiments, a mechanical driver (not shown) drives the distal end  22  of optical micro-probe  18  to execute scanning motions parallel and/or transverse to the axis of the GRIN fiber-size lens  26 . These scanning motions enable system  10  to collect optical data for two-dimensional or three-dimensional images of the sample  12 , i.e., a planar or full 3D image.  
         [0024]     Illumination beam  30  has a width that varies with distance from the end surface  28  of the GRIN fiber-size lens  26 . The beam width has a minimum value at an approximate focal point  34  of the GRIN fiber-size lens  26 , i.e., at a distance “f” from end face  28 . Typically, the distance “f” has a value from about 0.2 millimeters (mm) to about 1.5 mm, and exemplary values of “f” are greater than about 0.8 mm. The beam  30  has a divergence that is characterized by a rayleigh range “z”. Herein, the rayleigh range is half the length of the portion of the beam  30  that has a width less than about {square root}2 times the minimum width at the approximate focal point  34 . An exemplary GRIN fiber-size lens  26  has a rayleigh range greater than about 200 microns (μ), e.g., z≧300μ or 8 mm≧z≧300 μ.  
         [0025]     The focal distance and rayleigh range of GRIN fiber-size lens  26  depend on the radial profile of the refractive index in the GRIN lens and on the length of the GRIN lens. GRIN fiber-size lens  26  is either a conventional GRIN fiber-size lens or a new GRIN fiber-size lens with a gentler refractive index profile.  
         [0026]     Conventional GRIN fiber lenses are described in U.S. Pat. No. 4,701,011, which is incorporated herein by reference in its entirety.  FIG. 3A  shows the radial refractive index profile of one such GRIN fiber lens. The refractive index is constant over a range of values of the radius that correspond to the fiber&#39;s outer cladding and varies over values of the radius that correspond to the fiber&#39;s core. Restricting the refractive index variations to the core typically produces a GRIN fiber lens with a short focal length, less than about 0.7 mm, and a short rayleigh range, e.g., less than 200 μ.  
         [0027]      FIG. 3B  shows a radial refractive index profile of a new GRIN fiber-size lens  26  for which the profile&#39;s radial curvature is smaller in magnitude than in conventional GRIN fiber-size lenses. The smaller magnitude curvature causes the new GRIN fiber-size lens to have a longer focal length than the conventional GRIN fiber lens associated with the profile of  FIG. 3A . The new GRIN fiber-size lenses are described in co-pending U.S. patent application Ser. No. 09/896,789, filed Jun. 29, 2001, which is incorporated herein by reference in its entirety.  
         [0028]     In the profile of  FIG. 3B , the refractive index varies over the whole diameter of the lens. Thus, the new GRIN fiber-size lens has no outer cladding. The absence of cladding increases the radial range over which the refractive index varies permitting a larger optical mode, which results in the associated GRIN fiber-size lens having a longer rayleigh range than the GRIN fiber-size lens associated with the profile  FIG. 3A .  
         [0029]     Refractive index profiles are characterized by a parameter “g” that measures the radial curvature of the profile in the core of a GRIN fiber lens. In particular, the parameter g is defined as:  
       g   =         -     1     n   0         ⁢         ⅆ   2     ⁢     P   ⁡     (   r   )           ⅆ     r   2           ⁢     |     r   =   0             
 
 Here, “r” is radial distance for the axis of the GRIN fiber lens, n 0  is the value of the refractive index on the axis of the GRIN fiber lens, and P(r) is the value of the refractive index at the distance “r” from the axis of the fiber lens. 
 
         [0030]     Exemplary new GRIN fiber-size lenses have refractive index profiles whose radial curvatures are smaller in magnitude than those disclosed in Table 1 of “Analysis and Evaluation of Graded-Index Fiber-Lenses”, Journal of Lightwave Technology, Vol. LT-5, No. 9 (September 1987), pages 1156-1164, by W. L. Emkey et al, which is incorporated by reference herein in its entirety. The new GRIN fiber-size lenses  26  have a “g” that is less than 1.7×10 −6  μm −2 , preferable less than about 0.9×10 −6  μm −2  and more preferably less than about 5.0×10 −7  μm −2 . For 125 μm—diameter GRIN fiber lenses  18 , values of “g” are selected from the range 1.7×10 −6  μm −2  to 5.0×10 −7  μm −2  and preferably in the range 0.9×10 −6  μm −2  to 5.0×10 −7  μm −2  to provide good beam collimation.  
         [0031]     Referring again to  FIG. 2 , the above-disclosed refractive index profiles produce focal lengths and rayleigh ranges for GRIN fiber-size lens  26  that are consistent with the above-recited values. Some embodiments of optical micro-probe  18 ′ use a GRIN fiber-size lens  26  with a profile similar to that of  FIG. 3B , because such a profile provides a longer rayleigh range. The longer rayleigh range provides a larger usable depth range for sample probing. Typically, the usable depth of the optical micro-probe  18 ′ is about 1 to 8 rayleigh ranges from the focal point  34 .  
         [0032]      FIG. 1B  is a flow chart for a process  40  that uses system  10  of  FIGS. 1A and 2 . The process  40  includes positioning distal end  22  of the optical micro-probe to monitor a selected portion of sample  12  (step  42 ). The positioning includes selecting an orientation of the optical micro-probe  18  with respect to the sample surface  23  and selecting a lateral position and depth for the distal end  22  with respect to the sample surface  23 . After positioning the optical micro-probe  18 , source  14  transmits source light to the optical micro-probe  18  via splitter or circulator  16  (step  44 ). The transmitted source light passes through GRIN fiber-size lens  26 , which focuses the light into beam  30  (step  46 ). The region illuminated by the beam  30  produces the scattered or emitted light. The GRIN fiber-size lens  26  collects a portion of the light that is scattered or emitted by the illuminated region of the sample (step  48 ). The optical micro-probe  18  returns the light collected by the GRIN fiber-size lens  26  to the optical splitter or circulator  16 , which redirects a portion of the returned light to optical detector  20  (step  50 ). The detector  20  determines the scattering or emission characteristics of the region of the sample  12  from the light redirected thereto (step  52 ). Since the beam  30  has an intensity that varies with the beam width, the detector  20  primarily receives light from a region of the sample  12  that has a volume limited by the boundary of the beam  30 . The volume includes sample points within about 1 to 8 rayleigh ranges of focal point  34 . The light from the sample points and known position and orientation of optical micro-probe  18  enable using data from detector  20  to determine lateral positions and depths of the sample points backscattering or emitting light in some embodiments of system  10 .  
         [0000]     2. Interferometric Optical Monitoring and Imaging  
         [0033]     Various embodiments according to principles of the invention of this application include monitoring and imagining systems that determine depth and/or velocity data for a region of a sample that backscatters or emits light delivered by an optical micro-probe. Exemplary micro-probes include both probel  8 ′ with attached GRIN fiber-sized lens  26 , as shown in  FIG. 2 , and a single mode optical fiber without attached terminal GRIN fiber-sized lenses.  FIGS. 4A and 4B  show two such embodiments  60 ,  60 ′. To determine depths of sample regions, the systems  60 ,  60 ′ use “low-coherence interferometry” a method known to those of skill in the art.  
         [0034]     Each system  60 ,  60 ′ includes an interferometer with a measurement arm  62  and a reference arm  64 . The two arms  62 ,  64  receive light from a multi-chromatic source  66 , i.e., a low-temporal coherence source. Typically, source  66  is spatially coherent. The measurement arm  62  outputs light scattered by sample points in response to being illuminated by source light.  
         [0035]     Each system  60 ,  60 ′ interferometrically combines the light output by measurement arm  62  and reference arm  64 . The combined light provides an output signal sensitive to optical path differences between the two arms  62 ,  64 . Because of the low-coherence nature of source  66 , depth resolution is provided by the rapid fall off of the amplitude of the interference signal with increasing path length difference. Interferometric combining of light requires that the difference in the optical path lengths traversed by the light being combined, e.g., the path difference between the two arms  62 ,  64 , be less than the light&#39;s coherence length, e.g., the coherence length of the source  66 . Interference detector  74  uses the interferometrically combined light to determine one or more characteristics of the region of the sample  12  that produced scattered light, e.g., the intensity of the light scattered back into the optical micro-probe  18 . Thus, the sensitivity to optical path differences makes detector  74  sensitive to the depth of sample points  32  producing scattered light. The detector  74  is only sensitive to light produced by sample points  32  that are located within the sample depth range for which the optical path length difference between the measurement and reference arms  62 ,  64  is less than the coherence length of the source  66 .  
         [0036]     To increase depth resolution, a less coherent source  66 , e.g., a pulsed Ti-sapphire laser, is used in systems  60 ,  60 ′. The source  66  has a coherence length that is at least less than one centimeter and typically is less than one millimeter. In some embodiments, the source  66  has a coherence length that is as small as 100 microns or even 1 micron. Since interferometric combination only occurs if some optical path length differences between the measurement and reference arms  62 ,  64  are less than about one coherence length, this condition defines the depth resolution of the system  10 . For a sample depth resolution of 10 microns, the source  14  should produce an output beam that is only coherent for a time equal to about 10 −5  meters/{3×10 8  meters/second}=3×10 −14  seconds.  
         [0037]     The systems  60  and  60 ′ of  FIGS. 4A and 4B  include a Michelson interferometer and a Mach-Zehnder interferometer, respectively. Each system  60 ,  60 ′ has an optical splitter/combiner  68  that couples to one end of the measurement and reference arms  62 ,  64 . The optical splitter/combiner  68  transmits mutually coherent light from low-coherence source  66  to the measurement and reference arms  62 ,  64 . The measurement arm  62  includes optical micro-probe  18 . In the system  60 ′ of  FIG. 4B , the probe  18  connects to the measurement arm  62  through an optical circulator  65 . The probe  18  illuminates a sample region with source light and also collects light scattered produced by the illuminated sample region. In some embodiments, the optical micro-probe is a single-mode fiber  24  having a GRIN fiber-size lens  26  fused to its distal end  22 . The reference arm  64  includes a moveable reflector  76 , e.g. a moving mirror, and an acousto-optical modulator (AOM)  70 . The moveable reflector allows an operator to change the optical path length of the reference arm  64 , i.e., to scan different sample depths by moving the reflector  76 . The AOM  70  acoustically frequency shifts the source light received from the splitter/combiner  68  and enables velocities of sample points  34  to be measured (see below). Some embodiments include dispersion compensator  72  that corrects differences in chromatic dispersion or pulse broadening between light propagating in the measurement and reference arms  62 ,  64 . The construction of dispersion compensators is known to those of skill in the art.  
         [0038]     The interference detector  74  receives frequency-shifted light from the reference arm  64  and light scattered by the sample from the measurement arm  62 . The arms  62 ,  64  have optical path lengths that are equal to within about one coherence length of source  66  so that some light from the two arms  62 ,  64  interferometrically combines in the detector  74 , i.e., light produced by scattering at some sample depth. The detector  74  determines characteristics of regions of the sample producing light that interferometrically combines with light from the reference arm  64 . The moving reflector  76  enables an operator to adjust the optical path length difference between the reference and measurement arms  64 ,  62  so that sample depths can be scanned by the interference detector  74 . Through such scans, the systems  60 ,  60 ′ are able to generate images of the sample  12  as a function of sample depth.  
         [0039]      FIG. 4B  also shows an exemplary interference detector  74 . The exemplary interference detector  74  includes a 50/50 optical splitter/combiner  73  that produces signals with a 180° phase difference on its two output terminals. From the 50/50 optical splitter/combiner  73 , the 180° out of phase optical signals go to separate intensity detectors  75 . Outputs of the intensity detectors  73  couple to the inputs of a differential amplifier  77  whose output signal is representative of optical interference between signals from the reference and measurement arms  64 ,  62 .  
         [0040]     Referring to  FIG. 4C , AOM  70  includes a radio frequency (RF) source  78  and an optical medium  80 . The RF source  78  excites sound waves, i.e., phonons, in the optical medium  80 . The sound waves are directed along direction “Y” and have the source&#39;s RF frequency. A voltage oscillator  81  drives the RF source  78 . In some embodiments, the oscillator  81  is variable so that the phonon frequency is variable.  
         [0041]     Referring to  FIGS. 4A-4C , a photon in reference arm  64  may absorb or emit a phonon while propagating through the optical medium  80 . Absorption or emission of a phonon produces both a frequency-shift, i.e., +A/h, and a direction-change for the photon. Thus, the acoustically-driven medium  80  produces both directionally undeviated output light, i.e., photons that have neither absorbed nor emitted a phonon, and directionally deviated output light, i.e., photons that have absorbed or emitted a phonon. Momentum conservation fixes the directions of the deviated output light to be different from the direction of the undeviated output light. The AOM  70  is configured to deliver deviated output light of one frequency to the detector  74  and to not deliver the undeviated output light to the detector  74 . One embodiment screens out the undeviated light by imaging only deviated light, which has a new propagation direction, on an optical fiber that delivers light to moving reflector  72 . Thus, the interference detector  74  receives light whose frequency has been shifted by absorption or emission of a phonon in the AOM  70 .  
         [0042]     In some embodiments, the light makes two passes through the AOM  70 , and the AOM  70  screens out light whose frequency has not been shifted by the absorption or emission of two phonons. Then, this AOM  70  produces light whose frequency is shifted with respect to the optical source  66  by twice the frequency of the RF source  78 .  
         [0043]     Referring again to  FIGS. 4A-4B , the detector  74  obtains information on the displacement or velocity of the region of the sample  12  that backscatters source. The displacement or velocity information is encoded in the size of the Doppler shift caused by the velocity of the scattering region of the sample  12 . The AOM  70  enables detection of such Doppler shifts through phase-sensitive detection, which are known to those of skill in the art. In some embodiments, this detection technique enables a determination of both the sign and the magnitude of the velocities of scattering sample particles along the axis of optical micro-probe  18 . In other embodiments, this detection technique enables a determination of both the sign and the magnitude of displacements of scattering sample particles along the axis of optical micro-probe  18 .  
         [0044]     The AOM  70  provides light outputted by the reference arm  64  with a different frequency from the frequency of light outputted by the measurement arm  62 . In the absence of sample motion, this frequency difference is equal to the frequency of the RF energy driving the AOM  70 , i.e., if the reference arm  64  produces photons that absorb or emit one phonon. Sample motion at the depth for which the path difference between the measurement and reference arms  62 ,  64  vanishes changes the frequency difference between the light from the two arms  62 ,  64 , i.e., due to Doppler shifting. The detector  74  uses the magnitude of the change in the frequency difference between the light from the two arms  62 ,  64  to determine the speed of a sample particle producing scattering. The detector  74  uses the phase of the change in frequency difference, i.e., positive or negative, to determine the sign of the sample motion, i.e., towards or away from the optical micro-probe  18 . Standard electronic or optical techniques are known for determining both the magnitude and sign of the frequency difference between the light from the two arms  62 ,  64 .  
         [0045]     The systems  60 ,  60 ′ use the AOM  70  to determine information representative of velocities of sample points at a selected sample depth. Information representative of velocities of sample points includes signed displacements and velocities of the sample points along the axis of probe  18 . The systems  60 ,  60 ′ are also able to select different optical path lengths for the reference arm  64 , i.e., by moving reflector  76 . By scanning such optical path lengths, systems  60 ,  60 ′ are able to select different sample depths for which interferometric combination of scattered light from the measurement arm  62  and light from the reference arm  64  occurs. During such a scan, detector  74  determines sample region velocities as a function of distance from end  22  of optical micro-probe  18 , i.e., as a function of sample depth. This type of scan of sample velocities as a function of depth enables, e.g., for mapping blood flow rates in an artery of an animal or patient.  
         [0046]     The AOM  70  shifts light in the reference arm  64  by a single frequency. This simple form of the frequency shift enables the detector  74  to determine velocities in the sample  12 .  
         [0047]      FIG. 5  shows a medical diagnostic system  90  based on system  60  of  FIGS. 4A and 4C . The system  90  includes a differential amplifier  91  and an electronic filtering chain  92  that amplify and remove input noise, respectively. The system  90  also includes a multiplier  93  that combines a signal representative of the interferometrically combined optical signals from the measurement and reference arms  62 ,  64  with a signal representative of the RF signal driving RF source  78 . The output of the multiplier  93  goes to a fringe counter  94  that determines both the magnitude and sign of the velocity of a monitored portion of the sample. To determine the sign of the velocity, i.e., towards or away from optical micro-probe  18 , the counter  94  compares the signal from the multiplier  93  when the multiplier receives, i.e., via line  95 , different signals representative of the RF signal driving source  78 . The different signals are out of phase by 90°.  
         [0048]     The fringe counter  94  couples to a feed forward circuit  96  that in turn transmits information on the velocity and/or position of sample  12  to a controller  97 . The controller  97  is connected to a second diagnostic probe  98 , e.g., a monitoring electrode or a scanner for the same sample  12 . The controller  97  uses the information fed forward by circuit  96  to correct data that is output by the probe  98  for the effects of sample motion. In some embodiments, the controller  97  mechanically adjusts the position of the second diagnostic probe  98  to eliminate relative motion between the sample and probe  98 . In other embodiments, the controller  97  corrects the data collected by the second diagnostic probe  98  to compensate for the motion of the sample  12 , e.g., by displacing image scan data to eliminate motion induced smearing.  
         [0049]     Other embodiments of the invention will be apparent to those skilled in the art in light of the specification, drawings, and claims of this application.