Patent Application
20180299251
Optical Coherence Tomography (OCT) is an interference-based imaging technique. Known OCT systems and methods typically employ a reference arm and a sample arm. The sample arm delivers light to a sample to be imaged and collects light scattered or diffused from the sample. The scattered sample light is then mixed with light that is reflected from the reference arm (the “reference light”). If the sample light and the reference light are coherent, the mixing can produce an interference pattern that can be detected and then converted to an image. The coherence between the sample light and the reference light can be achieved by closely matching the path lengths of the sample and reference arms. Generally, the reference arm can be adjusted (e.g., changing the position of the reflector in the reference arm) to match the path lengths.
OCT can be useful in acquiring cross-sectional (tomographic) or volumetric images of a sample by scanning light across the sample. An image produced along the depth of the sample is conventionally termed as an A-scan. Each A-scan can provide information about the reflective or scattering properties of the sample as a function of depth at one position of the scanned beam. A cross-sectional image of the sample can be produced by combining neighboring A-scans. A volumetric image can be constructed from a group of B-scans, each of which can be an image of a planar slice into the sample. A B-scan, however, does not have to be a planar image. The B-scan can also be an image along a circle, and this cross-sectional view is then an annular scan around a point of interest in the sample.
Typically, at least three types of volumetric images are used in OCT imaging. A series of parallel B-scans can produce a rectangular, or raster volume scan; a series of B-scans at regular angular intervals can produce a radial volume scan; and an annular volume scan can be produced by a series of B-scans forming concentric rings. Each type of volumetric image can have its own advantages in particular circumstances. For example, rectangular volumes are frequently used in imaging the macula. Rectangular or annular scans are often used in the vicinity of the optic nerve head. Radial scans are often used in imaging the cornea.
OCT can have several desirable properties in imaging. First, the depth resolution, which can be dependent on precision of the depth scanning, can be independent from transverse resolution. High depth resolution can be achieved even at sites that may be not accessible by high numerical aperture (NA) beams, such as the fundus of the eye. A practical range of depth resolution can be on the order of 1 μm. Second, the interferometric technique used in OCT can provide high dynamic range and sensitivity (>100 dB), which can be beneficial in imaging of weakly scattering structures even in a scattering environment, thereby allowing “in situ optical biopsy.” Third, OCT is typically non-invasive and therefore can produce in vivo data without causing damages to the sample.
One issue with coherent imaging techniques, including OCT, however, is speckle noise. Speckle noise can be caused by the interference of light scattering from multiple points within a volume (or three-dimensional space) of the sample where light is focused and from which it is collected (this volume can be referred to as a resolution volume, a volumetric pixel or a voxel). More specifically, most surfaces, synthetic or natural, are rough on micro-scales (e.g., on the scale of the optical wavelengths). The rough surface (more specifically the reflectivity function of the surface) can be modeled as a collection of scatterers, each of which can scatter incident light. Because of the finite spatial resolution of an imaging system, at any time the light received by the detector can be regarded as being from a distribution of scatterers within the resolution volume. The scattered light adds coherently, i.e., the light from the scatterers interacts constructively and destructively depending on the relative phases of each scattered waveform. Constructive and destructive interference creates bright and dark dots in the image, thereby creating speckle noise, which can reduce the contrast of the resulting image thereby making boundaries between certain structures difficult to resolve.
Some known OCT systems and methods attempt to reduce speckle noise by employing incoherent averaging (also referred to as compounding) of several images (also referred to as snapshot). For example, averaging M images with uncorrelated speckle noise can reduce the speckle contrast by (M)1/2. The speckle contrast can be defined as the standard deviation of the noise divided by the mean intensity. Non-correlated speckle patterns can be obtained by various methods, including, but are not limited to, scanning from different angles, scanning several nearby regions, scanning with different incident wavelengths, and scanning with different polarizations. These methods can be referred to as angular, spatial, frequency, and polarization compounding, respectively. Compounding methods, however, normally compromise the resolution or depth of field of the resulting image when further reduction of speckle noise is pursued. Therefore, it can be challenging for compounding methods to eliminate speckle noise entirely.
Other known OCT systems and methods attempt to reduce speckle noise using image processing techniques, which can use adaptive filters and/or wavelet analysis to process the acquired images. These methods can reduce the appearance of noise. Such methods often, however, do not recover information that is lost or buried in the speckle.
Thus, a need exists for improved methods and devices for reducing speckle noise in OCT imaging.
Apparatus, systems, and methods for optical coherence imaging are described herein. In some embodiments, an apparatus includes a light splitter and a detector. The light splitter receives a spatially coherent light beam and directs a first portion of the spatially coherent light beam to a reference arm and a second portion of the spatially coherent light beam to a sample arm. The sample arm includes a phase scrambler at least partially in a path of the second portion of the spatially coherent light beam. The phase scrambler is configured to produce a sample light beam having a spatially variable phase. The sample arm also includes a controller, operably coupled to the diffuser, to change the spatially variable phase of the sample light beam. The detector is in optical communication with the reference arm and the sample arm, and is configured to detect an interference pattern produced by interference of the first portion of the spatially coherent light beam propagated through the reference arm and a scattered beam produced by scattering of the sample light beam by a sample propagated through the sample arm.
In some embodiments, an apparatus includes an optical arm of an optical coherence tomography system, a lens, a phase scrambler and a controller. The optical arm defines at least a portion of a light path, and is configured to be in optical communication with a light source that produces a spatially coherent light beam propagating along the light path. The lens is within the light path of the optical arm. The phase scrambler is disposed at least partially within the light path, and is configured to produce, from the spatially coherent light beam, a scrambled light beam having a spatially variable phase. The controller is operably coupled to the phase scrambler, and is configured to change the spatially variable phase of the scrambled light beam.
In other embodiments, a method includes transmitting a first portion of a spatially coherent light beam through a reference arm and transmitting a second portion of the spatially coherent light beam through a sample arm. The transmitting of the second portion includes A) changing a local phase of the second portion of the spatially coherent light beam to produce a sample light beam and B) transmitting the sample light beam toward a sample. The method further includes detecting an interference pattern produced by interference of the first portion of the spatially coherent light beam propagated through the reference arm and a scattered portion of the sample light beam scattered by and/or reflected from the sample via the sample arm.
In yet other embodiments, a method of coherence tomography includes transmitting a light beam to illuminate a resolution volume associated with a sample. The light beam is spatially modulated to introduce a first local phase change to a first portion of the light beam and to introduce a second local phase change to a second portion of the light beam, the second local phase change different than the first local phase change. The light beam is temporally modulated to produce a first speckle pattern at a first time in a first image associated with the resolution volume and to produce a second speckle pattern at a second time in a second image associated with the resolution volume. The second speckle pattern is different than the first speckle pattern. The method further includes averaging the first speckle pattern with the second speckle pattern to reduce speckle noise in a third image associated with the resolution volume.
In yet other embodiments, an apparatus includes a light source to produce a spatially coherent light, a light splitter in optical communication with the light source to split the spatially coherent light into a first beam and a second beam, a scanner in optical communication with the beam splitter to scan the second beam across at least a portion of a sample at a first speed so as to scatter and/or reflect light from the sample, and a detector in optical communication with the light splitter to detect interference between the first beam and the light scattered and/or reflected from the sample. The apparatus also includes a phase scrambler disposed within a Rayleigh range of an image plane of a lens to diffuse the second beam. An image of the sample at the image plane has a first magnification with respect to the sample. The apparatus further includes an actuator configured to move the diffuser in a direction substantially orthogonal to an optical axis of the diffuser at a second speed no less than a product of the first magnification and the first speed.
Apparatus, systems, and methods for optical coherence imaging are described herein. In some embodiments, an apparatus includes a light splitter and a detector. The light splitter receives a spatially coherent light beam and directs a first portion of the spatially coherent light beam to a reference arm and a second portion of the spatially coherent light beam to a sample arm. The sample arm includes a phase scrambler at least partially in a path of the second portion of the spatially coherent light beam. The phase scrambler is configured to produce a sample light beam having a spatially variable phase. The sample arm also includes a controller, operably coupled to the diffuser, to change the spatially variable phase of the sample light beam. The detector is in optical communication with the reference arm and the sample arm, and is configured to detect an interference pattern produced by interference of the first portion of the spatially coherent light beam propagated through the reference arm and a scattered beam produced by scattering of the sample light beam by a sample propagated through the sample arm.
In some embodiments, an apparatus includes an optical arm of an optical coherence tomography system, a lens, a phase scrambler and a controller. The optical arm defines at least a portion of a light path, and is configured to be in optical communication with a light source that produces a spatially coherent light beam propagating along the light path. The lens is within the light path of the optical arm. The phase scrambler is disposed at least partially within the light path, and is configured to produce, from the spatially coherent light beam, a scrambled light beam having a spatially variable phase. The controller is operably coupled to the phase scrambler, and is configured to change the spatially variable phase of the scrambled light beam.
In some embodiments, an apparatus includes a light source to produce a spatially coherent light, a light splitter, a scanner, a detector a phase scrambler and an actuator. The light splitter is in optical communication with the light source, and splits the spatially coherent light into a first beam and a second beam. The scanner is in optical communication with the light splitter, and is configured to scan the second beam across at least a portion of a sample at a first speed to produce a scattered beam scattered by the sample. The detector is in optical communication with the light splitter, and is configured to detect an interference between the first beam and the scattered beam. The phase scrambler is disposed within a Rayleigh range of an image plane of a lens, and is configured to modulate a local phase of the second beam. An image of the sample at the image plane has a first magnification with respect to the sample. The actuator is configured to move the phase scrambler in a direction substantially orthogonal to an optical axis of the phase scrambler at a second speed no less than a product of the first magnification and the first speed.
In some embodiments, a method of coherence tomography includes transmitting a light beam to illuminate a resolution volume associated with a sample. The light beam is spatially modulated to introduce a first local phase change to a first portion of the light beam and to introduce a second local phase change to a second portion of the light beam. The second local phase change is different than the first local phase change. The light beam is temporally modulated to produce a first speckle pattern at a first time in a first image associated with the resolution volume and to produce a second speckle pattern at a second time in a second image associated with the resolution volume. The second speckle pattern is different than the first speckle pattern. The method further includes averaging the first speckle pattern with the second speckle pattern to reduce speckle noise in a third image associated with the resolution volume.
In other embodiments, a method includes transmitting a first portion of a spatially coherent light beam through a reference arm and transmitting a second portion of the spatially coherent light beam through a sample arm. The transmitting of the second portion includes A) changing a local phase of the second portion of the spatially coherent light beam to produce a sample light beam and B) transmitting the sample light beam toward a sample. The method further includes detecting an interference pattern produced by interference of the first portion of the spatially coherent light beam propagated through the reference arm and a scattered portion of the sample light beam scattered by and/or reflected from the sample via the sample arm.
In yet other embodiments, a method of coherence tomography includes transmitting from a light source a light beam to a resolution volume associated with a sample. A first interference pattern is detected, at a first time and when the light beam is at a beam position relative to the sample. The first interference pattern is associated with the resolution volume, and is produced, in part, by a first scattered beam produced by scattering of the light beam from the resolution volume. The method includes changing a local phase of the light beam within the resolution volume of the sample. A second interference pattern is detected, at a second time after the changing and when the light beam is at the beam position relative to the sample. The second interference pattern is associated with the resolution volume, and is produced, in part, by a second scattered beam produced by scattering of the light beam having the changed local phase from the resolution volume. The first interference pattern and the second interference pattern are averaged.
In yet other embodiments, a method of coherence tomography includes transmitting from a light source a reference beam portion of a spatially coherent light beam to a reference member. The method includes, transmitting from the light source a sample beam portion of the spatially coherent light beam to a resolution volume associated with a sample. A local phase of at least one of the reference beam portion or the sample beam portion is changed. A first interference pattern is detected, at a first time and when the sample beam portion is in a beam position relative to the sample. The first interference pattern is associated with the resolution volume, and is produced based on the reference beam portion and the sample beam portion. The method includes changing, at a second time after the first time, the local phase of at least one of the reference beam portion or the sample beam portion. A second interference pattern is detected, at a third time and when the sample beam portion is in the beam position. The second interference pattern is associated with the resolution volume, and is produced based on the reference beam portion and the sample beam portion. The first interference pattern and the second interference pattern are averaged.
In yet other embodiments, an apparatus includes an elongated member, an optical transmission member, a lens and a phase scrambler. The elongated member is configured to be disposed within a bodily cavity, and defines a lumen. A side wall of the elongated member defines an opening. The optical transmission member is disposed within the lumen, and is configured to convey a sample light beam therethrough. The sample light beam is spatially coherent within the optical transmission member. The lens is disposed within the lumen and is optically coupled to the optical transmission member. The lens, the optical transmission member, and the opening of the elongated member define at least a portion of a sample light path through which the sample light beam is conveyed to a sample. The phase scrambler is disposed at least partially within a sample light path. The phase scrambler is configured to change a local phase of the spatially coherent sample light beam conveyed from the optical transmission member.
The term “about” when used in connection with a referenced numeric indication means the referenced numeric indication plus or minus up to 10 percent of that referenced numeric indication. For example, “about 100” means from 90 to 110.
In a similar manner, term “substantially” or “approximately” when used in connection with, for example, a geometric relationship, a numerical value, and/or a range is intended to convey that the geometric relationship (or the structures described thereby), the number, and/or the range so defined is nominally the recited geometric relationship, number, and/or range. For example, two structures described herein as being “substantially parallel” is intended to convey that, although a parallel geometric relationship is desirable, some non-parallelism can occur in a “substantially parallel” arrangement. By way of another example, a structured placed “approximately within an image plane” is intended to convey that, while the recited position is desirable, some tolerances can occur. Such tolerances can result from imperfections in optics that define the image plane, e.g., manufacturing tolerances, measurement tolerances, and/or other practical considerations. As described above, a suitable tolerance can be, for example, of ±10% of the stated geometric construction, numerical value, and/or range.
As used in this specification and the appended claims, the words “proximal” and “distal” refer to direction closer to and away from, respectively, an operator of the device. Thus, for example, the end of an imaging device adjacent or contacting the patient's body would be the distal end of the imaging device, while the end opposite the distal end would be the proximal end of the imaging device.
The indefinite articles “a” and “an,” as used herein in the specification and in the claims, unless clearly indicated to the contrary, should be understood to mean “at least one.”
The phrase “and/or,” as used herein in the specification and in the claims, should be understood to mean “either or both” of the elements so conjoined, i.e., elements that are conjunctively present in some cases and disjunctively present in other cases. Multiple elements listed with “and/or” should be construed in the same fashion, i.e., “one or more” of the elements so conjoined. Other elements may optionally be present other than the elements specifically identified by the “and/or” clause, whether related or unrelated to those elements specifically identified. Thus, as a non-limiting example, a reference to “A and/or B”, when used in conjunction with open-ended language such as “comprising” can refer, in one embodiment, to A only (optionally including elements other than B); in another embodiment, to B only (optionally including elements other than A); in yet another embodiment, to both A and B (optionally including other elements); etc.
As used herein in the specification and in the claims, the phrase “at least one,” in reference to a list of one or more elements, should be understood to mean at least one element selected from any one or more of the elements in the list of elements, but not necessarily including at least one of each and every element specifically listed within the list of elements and not excluding any combinations of elements in the list of elements. This definition also allows that elements may optionally be present other than the elements specifically identified within the list of elements to which the phrase “at least one” refers, whether related or unrelated to those elements specifically identified. Thus, as a non-limiting example, “at least one of A and B” (or, equivalently, “at least one of A or B,” or, equivalently “at least one of A and/or B”) can refer, in one embodiment, to at least one, optionally including more than one, A, with no B present (and optionally including elements other than B); in another embodiment, to at least one, optionally including more than one, B, with no A present (and optionally including elements other than A); in yet another embodiment, to at least one, optionally including more than one, A, and at least one, optionally including more than one, B (and optionally including other elements); etc.
The system 100 includes a light assembly 140, a beam splitter 110 (also generally referred to as a light splitter), a reference arm 120, and a sample arm 130. The light assembly 140 includes a light source 111 and a detector 142. The light source 111 can be any suitable light source of the types shown and described herein that produces a spatially coherent light beam 10 of a desired wavelength (or average wavelength, in instances where the light beam is a broadband beam). For example, in some embodiments, the light source 111 (and any of the light sources described herein) can be a super-luminescent diode (SLED or SLD). A SLD normally operates like edge-emitting laser diodes (EELD) but without optical feedback or a cavity. Super-luminescence may occur when the spontaneous emission experiences gain due to higher injection currents. The higher gain can cause a superlinear power increase and an increasing narrowing of the spectral width. The radiation emitted by an SLD can be amplified spontaneous emission (ASE) and have a low time-coherence. Since SLDs are generally implemented in wave-guide structures, the space-coherence of the emitted radiation can be accordingly high. The wavelength is determined by the material and its layering within the diode semiconductor. Practical wavelengths of SLD includes 675 nm, 820 nm, 930 nm, 1300 nm, 1550 nm, or any other wavelength known in the art.
In other embodiments, the light source 111 (and any of the light sources described herein) can include one or more of the following: a Ti:Sapphire laser at 800 nm, a Cr:forsterite laser at 1280 nm, a LED at 1240 nm-1300 nm, an amplified spontaneous emission (ASE) fiber source at 1300 nm-1550 nm, a super-fluorescence source such as a Yb-doped fiber (1064 nm), an Er-doped fiber (1550 nm), and a Tm-doped fober (1800 nm), a photonic crystal fiber at 725 nm or 1300 nm, or a thermal tungsten halogen source at 880 nm. The wavelength (or average wavelength) used in the SFOCT can be dependent on, for example, the desired penetration depth of the light in the sample.
The detector 142 and any of the detectors described herein can be any suitable detector that detects and/or receives light scattered and/or reflected from the sample and any reference elements. The detector 142 and any of the detectors described herein can include, for example, a charge coupled device (CCD). The detector 142 and any of the detectors described herein can also be referred to as a spectrometer. As described herein, the returned first portion 12R and returned second portion 14R of the light beam 10 are combined at the detector 142, which detects an interference pattern (when the system is properly aligned) produced by interference of the first return portion 12R and the second return portion 14R. An image of a sample 20 can then be extracted from the interference pattern detected by the detector 142.
The light splitter 110 receives a spatially coherent light beam 10 and splits the spatially coherent light beam into a first portion 12 and a second portion 14. The first portion 12 enters a reference arm 120, which further includes a dispersion compensation element 122 and a reference arm mirror 124. The dispersion compensation element 122 can compensate for dispersion introduced in the sample arm 130 to match the dispersion between the first portion of the light 12 and the second portion light 14 when they are returned to the detector 142 and combined to produce interferences. The reference arm mirror 124 can reflect the first portion 12 to propagate a first return portion 12R back to the light splitter 110. The light splitter the further reflects the first return portion 12R to a detector 142.
The second portion of the light 14 enters the sample arm 130. The sample arm 130 is configured to interact with a sample 20, such as a bodily tissue, to allow for optical coherence imaging of a portion of the sample 20. The sample arm 130 includes a phase scrambler 132 (also generally referred to as a diffuser or a local phase randomizer), a controller 134, and a series of lenses and/or optical components, as described below. Specifically, as shown in
The sample arm 130 also includes a first lens 131 having a first focal length f1, a second lens 133 having a second focal length f2, and a third lens 133 also having a second focal length f2. These three lens 131, 133, and 135 form a 4f configuration: the distance between the first lens 131 and the second lens 133 is f1+f2, the distance between the second lens 133 and the third lens 135 is 2×f2, and the distance between the third lens 135 and the sample 20 is f2. The 4f configuration can help relay the image of the sample 20 back to the detector 142.
The phase scrambler 132 is disposed at least partially in a light path 139 of the second portion 14 of the spatially coherent light beam 10. In this example, the phase scrambler 132 is a substantially transparent object (transparent to the wavelength of the light beam 10 used in the system 100), and includes a surface 138 (see the zoomed region Z1 of
The sample light beam, with the spatially variable phase, is then directed towards the sample 20, as shown in the zoomed region Z2 of
The controller 134 is operably coupled to the phase scrambler 132 to change the spatially variable phase of the second portion 14 of the spatially coherent light beam 10 as it passes through the phase scrambler 132. Similarly stated, the controller is operably coupled to the phase scrambler 132 and temporally changes the spatially variable phase of the sample light beam (i.e., the light beam that is propagated to the sample 20). The controller can be any suitable controller and/or can include any suitable mechanism to temporally change the phase of the sample light beam. For example, in some embodiments, the controller 134 includes an actuator to move the phase scrambler 132 within the path 139 of the spatially coherent light beam 10 in the phase scrambler 132. By moving the phase scrambler 132, the random distribution of micro diffusion centers on the surface 138 are moved, and thus the spatial variation of phase in the sample light beam is changed.
In some embodiments, the phase scrambler 132 is disposed approximately at an image plane of a lens (e.g., the lens 131 or the lens 133) within the sample arm 130. Similarly stated, in some embodiments, the phase scrambler 132 is aligned with a focal plane associated with the sample 20 and/or a lens (e.g., the lens 131 or the lens 133) within the sample arm 130. In some embodiments, the phase scrambler 132 is disposed within a Rayleigh range of the image plane associated with the sample 20 and/or a lens (e.g., the lens 131 or the lens 133) within the sample arm 130. In this manner, the phase scrambler 132 acts to change the local phase of the sample light beam, as described above. In some such embodiments, the controller 134 includes an actuator (not shown) to move the phase scrambler within the image plane, as shown by the arrow AA in
In some embodiments, the controller 134 includes an actuator (not shown) to rotate the phase scrambler within the image plane. The actuator can be, for example, a motor that rotates the phase scrambler 132 continuously during a sampling operation.
In some examples, the phase scrambler 132 and the controller 134 introduce spatial modulation into the light beam over the size of an imaging voxel (or a resolution volume of the sample 20). In some examples, the phase scrambler 132 and the controller 134 introduce temporal modulated over the course of A-scan acquisition (e.g., microseconds or longer). For example,
In use, the controller 134 and the detector 142 can be coordinated such that each image (or the original interference pattern) taken by the detector 142 includes a different speckle noise pattern created by a different diffusion introduced by the phase scrambler 132. In some examples, this can be achieved by tuning image taking rate of the detector 142 to be greater than the diffusion changing rate (or phase scrambling rate) of the controller 134 for the spatial light modulator. As described above, in some embodiments, the phase scrambler 132 has a static diffusion property (e.g., a ground glass), and the controller 134 can be configured to move the phase scrambler 132 by a substantial distance within the time interval of successive images taken by the detector 142. In these examples, the substantial distance can be, for example, comparable to the size of the second portion 14 of the spatially coherent light beam 10 (e.g., more than half of the diameter, more than a quarter of the diameter, more than a tenth of the diameter, etc.). In other embodiments, the distance moved by the phase scrambler 132 between successive images can be related to the wavelength (or average wavelength) of the light beam 10. For example, in some embodiments, the distance moved by the phase scrambler 132 between successive images can be about one wavelength (or average wavelength). In other embodiments, the distance moved by the phase scrambler 132 between successive images can be about two wavelengths (or two times the average wavelength).
In some embodiments, it is desirable to minimize movement when an image is being captured (to avoid blurring of the image), but maximize movement between the taking of images. In some embodiments, however, the phase scrambler 132 can be moved continuously during a sampling event (i.e., both during and between successive images). In some such embodiments, the phase scrambler 132 can be moved continuously during a sampling event at a speed of less than about one wavelength (or average wavelength). In some such embodiments, the phase scrambler 132 can be moved continuously during a sampling event at a speed of about one-third of a wavelength (or average wavelength).
The coordination between the controller 134 and the detector 142 can be carried out using any suitable software. For example, in some embodiments, the controller 134 can use software, such as Thorlabs APT (ThorLabs, Newton, N.J.).
In use, to reduce speckle noise in the images produced by the interference patterns, the system 100 can be operated according to an illustrative and non-limiting method below, as well as any other methods described herein. The first portion 12 of the spatially coherent light beam 10 is transmitted through the reference arm 120, and the second portion 14 of the spatially coherent light beam 10 is transmitted through the sample arm 130. In the sample arm 130, a time-varying local phase change is introduced into the second portion 14 of the spatially coherent light beam 10 by time-varyingly changing the diffusion produced by the phase scrambler 132. As described above, this can include moving the phase scrambler 132 along a direction perpendicular to the propagation direction of the second portion 14. After the phase scrambler 132, the sample light portion (shown with a non-planar wavefront) is then transmitted to the sample 20, where at least part of the sample light portion is reflected, diffused, and/or scattered back to the sample arm 130, which then transmits the reflected, diffused, or scattered part 14R to the detector 142. The detector 142 combines the second returned portion 14R with the first return portion 12R to produce an interference pattern. Multiple interference patterns can be taken by the detector 142. Each interference pattern is taken at a different timing moment. Due to the time-varying phase change introduced into the second portion 14 pf the light, the resulting interferences patterns taken at different timing moments include different and uncorrelated speckle noise patterns created by the diffusion. An image can be extracted from each interference pattern and the ultimate image of the sample can be produced by averaging the multiple images extracted from the multiple interference patterns.
Although the system 100 is shown and described as including a particular lens configuration (i.e., the 4f configuration), in other embodiments, any suitable lens configuration can be used in conjunction with a phase scrambler 132 (or any other phase scramblers shown herein) and to produce images according to any of the methods described herein. For example,
The phase scrambler 232 can be similar to the phase scrambler 132 shown and described above, and can located in any suitable position within the sample arm 230. For example, in some embodiments, the phase scrambler 232 can be located within an image plane, and can be configured to move (e.g., by a controller, not shown). In other embodiments, however, the phase scrambler 232 can be a stationary phase scrambler, of any type shown and described herein.
In some embodiments, a system can include any suitable structure and/or optical components to define the light paths through which the beams of light can be propagated. For example,
In other embodiments, depending on the operating wavelength of the imaging systems, other transmission devices, such as waveguides, can be used to transmit light or other radiation within the system.
Although the system 100 is shown and described as including a phase scrambler 132 within the sample arm 130, in other embodiments, an imaging system using time-varying phase scrambling can be implemented by placing a phase scrambler and/or a diffuser in the reference arm. This implementation can result in less aberration induced by the phase scrambler and/or the diffuser without reducing the optical power on the sample. This implementation includes focusing the light on the reference arm on to a phase scrambler, such as a moving diffuser, a rough mirror, or a spatial light modulator.
For example,
Although the optical systems described above, including the system 100, are shown and described as including a phase scrambler that is disposed within an image plane of a focused beam, in other embodiments, a system can include a phase scrambler disposed at any suitable location within a light path. Moreover, in some embodiments, systems and methods can include producing a local change of phase in a collimated (and not a focused) beam. Specifically, in some embodiments, a system can be configured employ phase scrambling in the Fourier domain. For example,
The imaging system 500 shown in
For example, the imaging system 700 shown in
In some embodiments, a Speckle Free Optical Coherence Tomography (SFOCT) system, such as the systems 100 and 200 shown above (or any other systems shown and described herein), can be constructed from commercially available OCT systems. For example, in some embodiments, a kit can include a phase scrambler of the types shown and described herein, a controller, and the necessary hardware to mount the phase scrambler and hardware within a commercial OCT system. In one example, a SFOCT system is built based on the Ganymede HR system manufactured by Thorlabs Inc. In this example, a diffuser can be placed at the focal plane of the original OCT probe and a new image plane can be projected by a 4f imaging system to visualize inside the sample. In this manner, the diffuser functions as a phase scrambler (similar to the phase scrambler 132 or an of the phase scramblers shown herein) to produce a local phase change in a light beam. Extension and addition of dispersion compensation elements to the reference arm can be employed to account for the addition of lenses and the extension of the sample arm.
In another example, a SFOCT system can be built based on the iFusion system manufactured by Optovue Inc. for retinal imaging and approved by the Food and Drug Administration (FDA). With this ophthalmic OCT, the image plane need not be replicated because an image plane is accessible inside the original OCT probe. The ophthalmic implementation of SFOCT can be simpler than that for other tissue imaging, because less change in the sample arm is made and the reference may not need any change at all (e.g., there is no need to include a dispersion compensation element).
For both examples using a commercial system, local and time-varying phase shifts can be implemented by placing a moving ground glass diffuser (a phase scrambler) at the OCT image plane. The diffuser can be placed, for example, inside a mount and moved by a motor in a plane substantially perpendicular to the optical axis (i.e., propagation direction of the light beams), as described above. Images can be acquired several times while the light beam is imaging the same location on the sample but propagating through different locations on the diffuser. In this manner, the time-varying pattern of the diffuser changes the speckle pattern of the image. After averaging several frames, speckle noise can be reduced significantly.
In some embodiments, SFOCT images in the above described systems, except for the human retina images, can be acquired using a Ganymede High-Resolution SD-OCT system (ThorLabs, Newton, N.J.) in accordance with any of the methods described herein. In some such embodiments, the light source can be a super-luminescent diode (SLED or SLD) operating at 30 kHz with a center wavelength at 900 nm and a full bandwidth of 200 nm (Δλ=800-1000 nm), which can provide 2.1 μm axial resolution in water. Further, the spectrometer can acquire 2048 samples for each A-scan. In some embodiments, at the beginning of each acquisition, the OCT can be programmed to measure the spectrum of the SLD for 25 times. These measurements can be used for the reconstruction of the OCT signal. The first lens of the imaging system can provide a lateral resolution of about 8 μm (FWHM) and depth of field (DOF) of about 143 μm in water (LSM03-BB, ThorLabs, Newton, N.J.). The 4f configuration can be implemented using lenses that provide a lateral resolution of about 4.2 μm (FWHM) and DOF of about 32 μm in water (LSM02-BB, ThorLabs, Newton, N.J.).
Due to the modification of the sample arm, including increasing the arm length, inclusion of lenses for the 4f configuration, and inclusion of the diffuser, dispersion is introduced. Accordingly, the reference arm can be extended by approximately 10 cm and dispersion compensation elements can be added (2×LSM02DC). The reference arm extension can be accomplished by, for example, placing metal extension rods between the OCT probe and the reference mirror.
In some embodiments, a retrofit of a commercially-available system can employ a diffuser to produce a local phase change in the light beam. Thus, the diffuser is a phase scrambler, as described herein. In such embodiments, the phase scrambler can be a ground glass diffuser with anti-reflection (AR) coating on one side (e.g., Thorlabs, DG10-1500-B and DG10-2000-B). In other embodiments, a phase scrambler can be a 3 μm lapped diffuser, which can be created by further lapping a 1500 grit diffuser. The diffuser can be mounted by a custom motorized mount with XYZ translation (e.g., based on CXYZ1, ThorLabs, Newton, N.J.). A conventional manual mount can also be used (e.g., ST1XY-S, ThorLabs, Newton, N.J.). The phase scrambler can be moved by the motors and controlled through computer software (e.g., Thorlabs APT). The movement of the phase scrambler can be perpendicular to the direction of the scan. For example, if the light beam is scanned along the X direction on the sample, the phase scrambler is then moved along the Y direction, and vice the versa. The diffuser can be moved back and forth at a speed of, for example, 0.3 mm/s with a range of 6.5 mm. The acceleration of the movement can be, for example, 1.5 mm/s/s.
In some embodiments, the phase scrambler is placed within the Rayleigh range of the Gaussian beam of the OCT system. In practice, the phase scrambler can be adjusted along the propagation direction of the light beams until a satisfactory image is acquired.
In some embodiments, images produced in SFOCT are normally averaged over several images taken at different timing moments. This averaging should not limit the application of SFOCT, at least because image averaging is already widely used in conventional OCT systems to reduce photon and thermal noise. In addition, as the acquisition rates if detectors increase, obtaining multiple frames can be completed within a shorter and shorter time, therefore without imposing any additional limitation to the application of SFOCT.
In some embodiments any of the optical systems, phase scrambler and methods described herein can be employed with an endoscopic imaging system. For example, in some embodiments, the free space optics (including the optical components, such as lenses, mirrors, transmission members or the like) described in any of the reference arms and/or sample arms described herein can be placed within an endoscope. For example,
The optical transmission member 860 can be any suitable optical component or structure to propagate light beams between a light source and/or a detector (not shown) and a sample (not shown) via the elongated member 850. Specifically, the optical transmission member 860 is disposed within the lumen 852, and conveys a sample light beam 10 therethrough. For example, in some embodiments, the optical transmission member 860 can be an optical fiber through which light can be propagated.
The sample light beam 10 can be produced by any suitable light source of the types shown and described herein, and is a spatially coherent light beam. The optical transmission member 860 also propagates a returned portion (not identified in
As shown, the system 800 includes a first lens 831, a second lens 833, and a third lens 835. The lenses, the optical transmission member 860, and the opening 853 of the elongated member 850 define at least a portion of a sample light path through which the sample light beam is conveyed to a sample (not shown). In some embodiments, the optical transmission member 860 can be an optical fiber coupled to the first lens 831 via a coupling member (or spacer) 861. The lenses can have any suitable configuration and/or arrangement within the elongated member 850. For example, in some embodiments the first lens 831 has a first focal length, the second lens 833 has a second focal length, and the third lens 833 also has a second focal length that matches that of the second lens. Thus, in such embodiments the three lens 831, 833, and 835 form a 4f configuration, as described above. In this manner, a conjugate image plane can be produced inside the elongated member 850 by using this lens arrangement. In other embodiments, however, the system 800 can employ any suitable lens configuration. Moreover, as shown, in some embodiments, an endoscopic system 800 can include a mirror 836 or other reflective element to propagate light through the opening 853 and to a sample.
The phase scrambler 832 is disposed at least partially in the sample light path, and reduces a spatial coherence of the spatially coherent light beam 10 that is propagated to the sample (e.g., via the opening 853). More particularly, as described herein, the phase scrambler 832 can produce a sample light beam having a spatially variable phase. In some embodiments, the phase scrambler 832 (and any of the phase scramblers and/or diffusers described herein) can include, but are not limited to, a ground glass element, a sandblasted glass element, an opal diffusing glass, or a holographic optical element, of the types shown and described herein.
In some embodiments, the phase scrambler 832 is disposed approximately at an image plane of a lens (e.g., the lens 831 or the lens 833) within the elongated member 850. Similarly stated, in some embodiments, the phase scrambler 832 is aligned with a focal plane associated with the sample and/or a lens (e.g., the lens 831 or the lens 833) within the elongated member 850. In some embodiments, the phase scrambler 832 is disposed within a Rayleigh range of the image plane associated with the sample and/or a lens (e.g., the lens 831 or the lens 833) within the elongated member 850. In this manner, the phase scrambler 832 acts to change the local phase of the sample light beam, as described above.
In some such embodiments, the system 800 includes an actuator 834 (or controller) to move the phase scrambler 832 within the image plane. In some such embodiments, the actuator 834 includes an actuator (not shown) to move the phase scrambler along a direction non-parallel (e.g., substantially perpendicular to) a propagation direction of the second portion of the spatially coherent light beam 10. In some embodiments, the actuator 834 is configured to rotate the phase scrambler 832 within the image plane. The actuator can be, for example, a motor that rotates the phase scrambler 832 continuously during a sampling operation. In other embodiments, however, the phase scrambler 832 can be at a fixed position within the elongated member 850, as described herein.
Although shown as including three lenses, in other embodiments, an endoscopic SFOCT system can include any suitable lens configuration. For example,
The optical transmission member 960 can be any suitable optical component or structure to propagate light beams between a light source and/or a detector (not shown) and a sample (not shown) via the elongated member 950. Specifically, the optical transmission member 960 is disposed within the lumen 952, and conveys a sample light beam 10 therethrough. For example, in some embodiments, the optical transmission member 960 can be an optical fiber through which light can be propagated.
As shown, the system 900 includes a single lens 931. The lens 931, the optical transmission member 960, and the opening 953 of the elongated member 950 define at least a portion of a sample light path through which the sample light beam is conveyed to a sample (not shown). Moreover, as shown, in some embodiments, an endoscopic system 900 can include a mirror 936 or other reflective element to propagate light through the opening 953 and to a sample.
The phase scrambler 932 is disposed at least partially in the sample light path, and reduces a spatial coherence of the spatially coherent light beam 10 that is propagated to the sample (e.g., via the opening 953). More particularly, as described herein, the phase scrambler 932 can produce a sample light beam having a spatially variable phase. In some embodiments, the phase scrambler 932 (and any of the phase scramblers and/or diffusers described herein) can include, but are not limited to, a ground glass element, a sandblasted glass element, an opal diffusing glass, or a holographic optical element, of the types shown and described herein. As shown, the phase scrambler 932 is disposed at the tip (or end surface) of the optical transmission member 960. Specifically, the phase scrambler 932 is disposed between a coupling member (or spacer) 961 and the optical transmission member 960.
As described above, in some embodiments any of the systems and methods described herein can include a movable phase scrambler that introduces a time-varying shift can into illuminating light beams in imaging systems. In some embodiments, any of the phase scrambler can be a transmissive optical member, such as a diffuser. To improve speckle reduction, the random phases introduced by the phase scrambler (or diffuser) can be evenly distributed between 0 to 2π at the beam waist. Stated differently, the surface height variation of the diffuser can be λ/Δn, where λ is the operating wavelength of the imaging system and Δn is the difference of refractive index between the diffuser and air. For example, to obtain this phase shift using a diffuser made of glass with a refractive index of 1.5 (NBK-7) and light sources with a center wavelength of 900 nm, the total thickness variation of the diffuser can span at least 1.8 μm. On the other hand, deflection of light by the diffuser, which is more probable in a ground glass diffuser with a wide thickness range, may reduce the OCT signal.
Thus, the amount of surface roughness of the transmissive phase scrambler (or diffuser) should be sufficient to introduce the desired local phase shift, while also minimizing power loss as the beam is propagated through the diffuser. Moreover, if the surface roughness (i.e., the peak-to-peak variation in the surface structures) exceeds more than about 5 microns, the resulting images become blurry. Said another way, if the variation in the surface finish is too great, the light beam will lose temporal coherence, and thus the axial resolution will be limited.
To evaluate different diffusers for use as phase scramblers, three types of diffusers are used and characterized with respect to their thickness and roughness using a 3D optical profiler and an atomic force microscopy (AFM). The first diffuser (also the roughest diffuser) is an off-the-shelf 1500 grit diffuser with AR coating (Thorlabs Inc.). The second diffuser is a custom made 2000 grit diffuser (Thorlabs Inc.). The third diffuser is made by further grinding (lapping) the 1500 diffuser with 3 μm particles.
As discussed above, implementing SFOCT with diffusers having a different surface profile can have different effects on the optical power on the sample, the signal (also referred to as the OCT signal), and the lateral resolution. For example,
Although the roughest diffuser (1500 grit) may reduce the OCT signal and the lateral resolution the most (35.3% increase in the size of the point spread function, versus 6.2% in the 2000 grit diffuser), it can achieve the best overall performance in terms of speckle removal and appearance of small detail in tissue. This tradeoff may be avoided by careful design and fabrication of a designated diffuser.
Although the phase scramblers described herein can include a diffuser of the types shown and described herein, and can be moved relative to a light path to introduce a local phase difference that is changed between image samples, in other embodiments, any suitable phase scrambler can be used in any of the systems and methods described herein. For example, in some embodiments, any of the systems and methods described herein can include a phase scrambler that is at a fixed location. In other embodiments, any of the systems and methods described herein can include a phase scrambler does not transmit light therethrough. For example, in some embodiments, a phase scrambler, such as the phase scrambler 132, includes a spatial light modulator configured to change the diffusion of a portion (i.e., a split beam) of spatially coherent light via at least one of a mechanical force, an electrical field, a magnetic field, or a thermal field.
Various types of spatial light modulators can be used as a phase scrambler. In one example, the spatial light modulator can be an electrically addressed spatial light modulator (EASLM). The diffusion in an electrically addressed spatial light modulator can be created and changed electronically (similar to most electronic displays). EASLMs usually receive input via a conventional interface such as Digital Visual Interface (DVI) or Video Graphics Array (VGA) input. An example of an EASLM is the Digital Micromirror Device using ferroelectric liquid crystals (FLCoS) or nematic liquid crystals (Electrically Controlled Birefringence effect). In another example, the spatial light modulator can be an optically addressed spatial light modulator (OASLM). The diffusion in an optically addressed spatial light modulator, also known as a light valve, can be created and changed by shining light encoded with an image on the front or back surface of the OASLM. A photosensor can be employed to allow the OASLM to sense the brightness of each pixel and replicate the pattern in the encoded light using liquid crystals. Typically, as long as the OASLM is powered, the diffusion pattern is retained even after the light is extinguished. An electrical signal can be used to clear the whole OASLM at once.
Effective Resolution of SFOCT
One advantage of SFOCT using the systems and methods described herein can be the improved effective resolution so that closely-spaced scatters or other features can be distinguished (resolved). This improved effective resolution can be demonstrated and quantified by imaging an infinitesimally small gap. For example,
More specifically, four 5 mL agarose phantoms embedded with various scattering agents can be created using a stock solution of 1% agarose in water. The agarose solution can be prepared on a hot plate with a magnetic stirrer and kept at a constant temperature of 60° C. to prevent curing or clumping. Three different scattering agents were employed: 0.3-1 μm TiO2 rutile powder (e.g., Atlantic Equipment Engineers, Upper Saddle River, N.J.), 21 nm (primary particle size) TiO2 anatase nanopowder (e.g., Sigma Aldrich Co. St. Louis, Mo.), and OD 500 gold nano-rods (GNRs) with peak absorption at 745 nm. Two phantoms were made with a low and a high concentration of GNR, respectively. The high concentration phantom included 100 μL of GNR for every 5 mL of base, and the low concentration used 50 μL of GNR for every 5 mL of base. For phantoms with TiO2 as the scattering agent, 0.009 grams of the TiO2 were ultrasonically dispersed in 1 mL Millipore water using a water bath sonicator to prevent aggregation. The solution can be sonicated for four 30 second intervals with a two-minute gap between each round to prevent overheating. Scattering agents can be slowly added to 5 mL of uncured agarose at 60° C. with continuous stirring. The final solution can be stirred for one minute before being poured into 5 mL plastic petri dishes. Two hours or longer curing were carried out before being used in SFOCT for imaging.
Although the lateral resolution of SFOCT can be lower compared to that in OCT when measured on a glass test chart (e.g., shown in
Pixel Value Statistics in SFOCT
OCT speckle noise normally follows a Rayleigh distribution. This can be problematic since the Rayleigh statistics may dominate the image and obscure real features within the sample. SFOCT according to the embodiments described herein can reduce speckle noise by shifting the pixel value statistics from a Rayleigh distribution toward the expected distribution of scatters in a sample.
Speckle contrast can be theoretically and experimentally reduced by (M)1/2 using SFOCT conducted by the systems and according to the methods described herein. Since each image can be acquired at a similar angle, sample position, and illumination wavelength, increasing the number of compounded images normally does not correlate to an inherent degradation in resolution. Because increasing the number of uncorrelated images does not reduce resolution, it is possible to create many images and subsequently eliminate speckle noise entirely. An approximate mathematical description of this phenomenon is given by:
In which I is the pixel value after averaging M scans obtained at different times and with different local phase shifts. N is the number of scatters inside a voxel, with scattering amplitudes an for the nth scatterer and locations that cause a relative phase shift of φn. θnm is the local phase shift in the location of scatterer n, and which changes in time. In conventional OCT θnm is constantly equal to zero; in SFOCT, however, it is a random variable with a uniform distribution between 0 and 2π.
The shift of statistics can be experimentally demonstrated by measuring pixel value statistics of a phantom made of gold nanorods (GNRs) disposed in an agarose gel. Similar to the phantoms made with TiO2, owing to the high backscattering of the metallic particles and their high concentration, these phantoms can be useful models for turbid media and produce speckle statistics as expected for conventional OCT imaging.
The pixel value statistics obtained with SFOCT resemble a Poisson distribution in which each event contributes a value that is equal to the backscattering of a single GNR. This distribution more closely matches the expected distribution of GNRs randomly dispersed within the phantom. As predicted by equation (1), increasing the number of averages reduces speckle noise and thus reduces the broadness of the distribution of pixel values. Because the phantoms are composed of a random spatial distribution of particles, a uniform signal is not expected when speckle noise is eliminated. Because SFOCT is capable of eliminating speckle, it can more closely approximate the actual distribution of scatterers in the sample.
As mentioned above, conventional OCT speckle noise follows a Rayleigh distribution. SFOCT reduces speckle and shifts the pixel value statistics from a Rayleigh distribution toward the expected distribution of scatterers in a sample, which is a Poisson distribution in which each event contributes a value that is equal to the backscattering of a single GNR. The two expressions for the two distributions are:
In which I is the pixel value and I
the average pixel intensity. k is the number of particles in a voxel, which can be assumed to contribute equally to the OCT signal and λp is the average number of particles in a voxel. To obtain λp the histogram of the pixel values is fit to a Poisson distribution. From that, the contribution of each equivalent particle to the OCT signal is derived as Ip=
I
/λp.
To further validate that SFOCT conducted using the systems and in accordance with the methods described herein removes speckle, experimental data can be compared to the √{square root over (M)} model. In the case of an inherent signal variation, the reduction in the normalized standard deviation (STD), can be described by:
In which σ is the measured STD in a region of interest, I
the average pixel intensity and C is the normalized STD which includes the speckle contrast and the inherent signal variation of the sample. σ2 is the variance in pixel values, σ02 is the intrinsic variation in signal caused only by the variation in number of particles in a voxel, and σspeckle2 is the variance of the speckle noise, which reduces by a factor of M during compounding.
Equations (4a)-(4c) show that the normalized speckle, as defined below, should reduce by √{square root over (M)}:
On a logarithmic:
Imaging Fine Structures Using SFOCT
The speckle reduction achieved with SFOCT, performed by any of the systems and in accordance with any of the methods described herein, can reveal fine structures that are otherwise obscured by noise. The capability of SFOCT can be demonstrated by imaging polystyrene beads with 3 μm diameter embedded inside a GNR and agarose phantom.
As shown in
Biomedical Imaging of SFOCT
One of the biomedical advantages of OCT is that it can provide noninvasive high resolution images inside living tissues. Strong speckle artifacts, however, drastically limit insight regarding fine anatomical structures. These limitations become obvious upon comparison with histological tissue sections. By removing the significant contribution of speckle noise, SFOCT is capable of rendering in vivo images that approach histological detail.
As one example,
As seen from
As also evident by the images and characterization presented in
Another common clinical application of OCT is for ophthalmic imaging.
In
Using SFOCT, the lamella structure of the stroma along with clear boundaries between the other layers of the cornea can be observable. Due to speckle noise, conventional OCT is not able to show clear boundaries between layers or the structure of the stroma.
In the retinal systems, the power of the light source is normally limited to a certain safety level (e.g., due to ANSI safety guidelines). A finer grit diffuser produced by lapping a 1500 grit diffuser with 3 μm particles is used as the phase scrambler in the SFOCT system for this imaging. As discussed above, the finer diffuser can produce images with a higher signal to noise ratio but may also reduce less speckle noise. Although speckle noise may still be present in the retina images, the retinal layers are better defined in SFOCT when compared to conventional OCT.
Other than ophthalmic applications, OCT is gaining popularity in dermatology owing to its potential for doing noninvasive biopsy. Speckle noise, however, may prevent seeing clear boundaries between anatomical objects and limits the visibility and identification of small or low contrast structures. Speckle elimination can enhance the diagnostic capabilities of OCT.
As seen from
Other Speckle Noise Processing Techniques
Speckle noise may be reduced by other techniques, such as spatial compounding, adaptive Wiener filtering, symmetric nearest-neighbor, and hybrid median filter. These cases can be used in combination with SFOCT techniques. For example,
Techniques that are disclosed in this application use optical coherence tomography (OCT) as an illustrating and non-limiting example. Techniques can be implemented in all types of OCT, including time-domain OCT, swept source OCT, and spectral domain OCT. In addition, techniques disclosed herein can also be applied in any other coherence imaging techniques, such as holography, interference based profilometer, and coherent imaging and microscopy.
Additional Applications
The reduction of speckle noise performed in accordance with the systems and methods described herein not only reduces noise in visualizing the structure of the sample, but also reduces noise when analyzing the spectral characteristics of a sample when performing spectral analysis. Spectral analysis is performed to visualize the spectrum of scattered light from a sample, in addition to the location of the scatterer, and can reveal, for example, amount of blood oxygenation and a location of a contrast agent. Speckle noise, which is caused by interference between coherently scattered light, has spectral components therein. Thus, speckle noise is wavelength dependent, and therefore, it creates noise in the spectral analysis of the sample.
To evaluate the impact of the SFOCT systems and methods described herein on spectral analysis, a sample as imaged using both conventional OCT and SFOCT in accordance with the methods described herein, and spectral analysis was then performed. The spectral analysis can be performed using any suitable method, such as, for example, the method disclosed in “Contrast-enhanced optical coherence tomography with picomolar sensitivity for functional in vivo imaging,” by O. Liba et al., Sci. Rep., vol. 6, p. 23337, March 2016. The results of the spectral analysis for each frame were then averaged. Each frame has a different spectral-speckle pattern because of the different phases projected by the phase scrambler (or diffuser). After averaging, the spectral-speckle noise is considerably reduced. The results are shown in
In some embodiments, a method includes transmitting a light beam to illuminate a resolution volume associated with a sample. The light beam is spatially modulating to introduce a first local phase change to a first portion of the light beam and to introduce a second local phase change to a second portion of the light beam. The second local phase change different than the first local phase change. The light beam is then temporally modulating between successive image capture events to produce a first speckle pattern at a first time in a first image associated with the resolution volume and to produce a second speckle pattern at a second time in a second image associated with the resolution volume. The second speckle pattern is different than the first speckle pattern. A spectral analysis is then performed on a series of images, including those with the first speckle pattern with the second speckle pattern. The first speckle pattern and the second speckle pattern are averaged to reduce speckle noise in a third image associated with the resolution volume. The third image can be, for example, associated with a spectral analysis.
In some embodiments, the SFOCT apparatus and method described herein can be applied to the measurement of an absorption profile of a sample. In such embodiments, the reduction and/or elimination of the speckle noise produces a more accurate measurement of the absorption profile than measurements taken using convention OCT methods. In particular, studies have shown that using OCT to measure the local attenuation coefficient in tissue can provide diagnostic information, such as the detection of tumor margins. The attenuation coefficient of a sample may be calculated by fitting the optical coherence tomography signal intensity to a function that includes the effects of the Beer-Lambert law (an exponential function), the confocal function, OCT roll-off and multiple scattering.
In some embodiments, a method includes applying SFOCT to the measurement of local attenuation coefficients. In some embodiments, a method includes varying a local phase of a sample light beam, as described herein, between successive images to reduce speckle noise. The method further includes evaluating the signal intensity of the images to determine a boundary of a structure within the sample. As described herein, by applying the SFOCT methods described herein to remove and/or reduce the speckle noise, a more precise fit can be produced.
To compare the precision of the fit between conventional OCT and SFOCT, an exponential fit to the signal intensity was performed based on scans conducted using both techniques. The simplified model described herein assumes that tissue attenuation, governed by the Beer-Lambert law, is the most dominant factor in the decrease of signal intensity as a function of depth. The precision of the fit between images of a fingertip produced using conventional OCT and SFOCT according to the methods described herein was compared. (see e.g.,
The Beer-Lambert law is given by the following expression:
i(z)∝√{square root over (exp(−2μz))}=exp(−bz) (7)
in which i (z) is the depth dependent OCT or SFOCT signal intensity and μ is the attenuation coefficient. b is the exponential coefficient that we are attempting to find by fitting the signal to an exponential function.
In some embodiments, a method can include changing the local phase of a light beam during optical coherence tomography to reduce and/or eliminate speckle noise to improve the automated segmentation of structures in the imaged tissue. For example, when analyzing the retina, segmentation is used to determine the thickness of the retinal layers and diagnose numerous conditions. In some embodiments, reducing speckle noise using the SFOCT methods and systems described herein can improve segmentation results.
To evaluate the effects of SFOCT on segmentation algorithms, segmentation of mouse retina was used as a test case. The external limiting membrane (ELM) was analyzed, because this thin layer is known to be one of the most difficult layers to detect in retina imaging using conventional OCT methods. Particularly, imaging of this thin layer often results in poor signal-to-noise ratio (SNR). Many known retina layer segmentation algorithms use graph cut segmentation. Interestingly, a recent study examined the effect of SNR of the image on segmentation error rate, and found that at a border case in which the SNR is 2, the expected error rate would be 15.8%. In many instances, this level of error is considered as practically sufficient for many segmentation problems. Of course, the higher the SNR, the lower the resulting error rate. Thus, increasing the SNR using the methods and systems described herein can provide improved segmentation.
To test this method, a mouse retina was imaged using convention OCT methods and SFOCT methods according to an embodiment. A region of interest (ROI) was selected around the ELM (See,
Where Ip is the image log intensity for pixel p balanced such that each column has the same mean intensity, FG are pixels within the foreground area, BG are pixels within the background area.
The signal to noise ratios were measured as being 0.76 for the OCT image, 2.56 for OCT with 24 μm lateral smoothing, and 3.30 for SFOCT with no lateral smoothing. These measurements imply that it is possible to reach similar levels of SNR (and similar segmentation quality as a result) with both OCT and SFOCT only if a significant lateral smoothing is applied to the OCT image (
We further quantified SFOCT improvements to segmentation quality. One common way to measure segmentation quality is to compute the mean absolute difference (MAD) between an automated algorithm and a manual segmentation (human eye) in comparison to the MAD of two manual segmentations. In other words, MAD between two manual segmentations is considered as a reference point when measuring segmentation quality. Therefore, to assess the possible improvement in segmentation quality resulting from SFOCT, we asked 3 study-blinded, unbiased subjects to segment the ELM within the given ROI. The subjects returned mean MAD of 0.674 pixels for OCT images, 0.339 pixels for OCT with lateral smoothing, and 0.341 pixels for SFOCT. The significant difference between segmentation of OCT and SFOCT images (t-test p≤0.02) provides evidence that, in this case, a future automated SFOCT segmentation could be significantly more accurate than an OCT image segmentation algorithm. No significant difference in MAD was observed when comparing OCT with lateral smoothing and SFOCT
Additional Methods
As described above, the systems and methods described herein can be used in any suitable application. Moreover, any of the above-described applications can employ any of the methods of scanning, average and phase scrambling described herein. For example, any of the systems and applications described herein can be performed using the method of optical coherence tomography shown by the flow chart in
The light beam is temporally modulated to produce a first speckle pattern at a first time in a first image associated with the resolution volume and to produce a second speckle pattern at a second time in a second image associated with the resolution volume, at 53. The second speckle pattern is different than the first speckle pattern. The temporal modulation can be performed using any suitable device or system described herein, such as any of the phase scramblers described herein. For example, in some embodiments, the temporal modulation can be performed by moving a phase scrambler within a light path between successive image capture events.
The method further includes averaging the first speckle pattern with the second speckle pattern to reduce speckle noise in a third image associated with the resolution volume, at 54.
As another example, any of the systems and applications described herein can be performed using the method of optical coherence tomography shown by the flow chart in
A local phase of the light beam within the resolution volume of the sample is changed, at 63. The local phase change can be produced using any suitable device or system described herein, such as any of the phase scramblers described herein. For example, in some embodiments, the spatial modulation can be performed using a ground glass diffuser, as described above.
The method further includes detecting, at a second time after the changing and when the light beam is at the beam position relative to the sample, a second interference pattern associated with the resolution volume, at 64. The second interference pattern is produced, in part, by a second scattered beam produced by scattering of the light beam having the changed local phase from the resolution volume. The method further includes averaging the first interference pattern and the second interference pattern, at 65.
While various inventive embodiments have been described and illustrated herein, a variety of other means and/or structures for performing the function and/or obtaining the results and/or one or more of the advantages described herein. More generally, all parameters, dimensions, materials, and configurations described herein are meant to be examples and that the actual parameters, dimensions, materials, and/or configurations will depend upon the specific application or applications for which the embodiment(s) is/are used. Many equivalents to the specific embodiments described herein are possible. It is, therefore, to be understood that the foregoing embodiments are presented by way of example only and that, within the scope of the appended claims and equivalents thereto, embodiments may be practiced otherwise than as specifically described and claimed. Embodiments of the present disclosure are directed to each individual feature, system, article, material, kit, and/or method described herein. In addition, any combination of two or more such features, systems, articles, materials, kits, and/or methods, if such features, systems, articles, materials, kits, and/or methods are not mutually inconsistent, is included within the scope of the present disclosure.
The above-described embodiments can be implemented in any of numerous ways. For example, embodiments of designing and making the technology disclosed herein may be implemented using hardware, software or a combination thereof. When implemented in software, the software code can be executed on any suitable processor or collection of processors, whether provided in a single computer or distributed among multiple computers.
Further, it should be appreciated that a computer may be embodied in any of a number of forms, such as a rack-mounted computer, a desktop computer, a laptop computer, or a tablet computer. Additionally, a computer may be embedded in a device not generally regarded as a computer but with suitable processing capabilities, including a Personal Digital Assistant (PDA), a smart phone or any other suitable portable or fixed electronic device.
Also, a computer may have one or more input and output devices. These devices can be used, among other things, to present a user interface. Examples of output devices that can be used to provide a user interface include printers or display screens for visual presentation of output and speakers or other sound generating devices for audible presentation of output. Examples of input devices that can be used for a user interface include keyboards, and pointing devices, such as mice, touch pads, and digitizing tablets. As another example, a computer may receive input information through speech recognition or in other audible format.
Such computers may be interconnected by one or more networks in any suitable form, including a local area network or a wide area network, such as an enterprise network, and intelligent network (IN) or the Internet. Such networks may be based on any suitable technology and may operate according to any suitable protocol and may include wireless networks, wired networks or fiber optic networks.
The various methods or processes (outlined herein) may be coded as software that is executable on one or more processors that employ any one of a variety of operating systems or platforms. Additionally, such software may be written using any of a number of suitable programming languages and/or programming or scripting tools, and also may be compiled as executable machine language code or intermediate code that is executed on a framework or virtual machine.
In this respect, various disclosed concepts may be embodied as a computer readable storage medium (or multiple computer readable storage media) (e.g., a computer memory, one or more floppy discs, compact discs, optical discs, magnetic tapes, flash memories, circuit configurations in Field Programmable Gate Arrays or other semiconductor devices, or other non-transitory medium or tangible computer storage medium) encoded with one or more programs that, when executed on one or more computers or other processors, perform methods that implement the various embodiments of the disclosure. The computer readable medium or media can be transportable, such that the program or programs stored thereon can be loaded onto one or more different computers or other processors to implement various aspects of the disclosure.
The terms “program” or “software” are used herein in a generic sense to refer to any type of computer code or set of computer-executable instructions that can be employed to program a computer or other processor to implement various aspects of embodiments as discussed above. Additionally, it should be appreciated that according to one aspect, one or more computer programs that when executed perform methods of the disclosure need not reside on a single computer or processor, but may be distributed in a modular fashion amongst a number of different computers or processors to implement various aspects of the disclosure.
Computer-executable instructions may be in many forms, such as program modules, executed by one or more computers or other devices. Generally, program modules include routines, programs, objects, components, data structures, etc. that perform particular tasks or implement particular abstract data types. Typically, the functionality of the program modules may be combined or distributed as desired in various embodiments.
Also, data structures may be stored in computer-readable media in any suitable form. For simplicity of illustration, data structures may be shown to have fields that are related through location in the data structure. Such relationships may likewise be achieved by assigning storage for the fields with locations in a computer-readable medium that convey relationship between the fields. Any suitable mechanism, however, may be used to establish a relationship between information in fields of a data structure, including through the use of pointers, tags or other mechanisms that establish relationship between data elements.
Also, various disclosed concepts may be embodied as one or more methods, of which an example has been provided. The acts performed as part of the method may be ordered in any suitable way. Accordingly, embodiments may be constructed in which acts are performed in an order different than illustrated, which may include performing some acts simultaneously, even though shown as sequential acts in illustrative embodiments.
This application benefit of priority to U.S. Provisional Application Ser. No. 62/243,466, entitled “Methods and Apparatus for Speckle-Free Optical Coherence Imaging,” filed Oct. 19, 2015, which is incorporated herein by reference in its entirety.
This invention was made with Government support under contracts NSF 1438340 awarded by National Science Foundation, CA 151459 awarded by the National Cancer Institute and OD 012179 awarded by the National Institutes of Health. The Government has certain rights in the invention.
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/US16/57656 | 10/19/2016 | WO | 00 |
| Number | Date | Country | |
|---|---|---|---|
| 62243466 | Oct 2015 | US |
