US20190150744A1 - Pulsed ultrasound modulated optical tomography with increased optical/ultrasound pulse ratio - Google Patents
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Abstract
Description
- Pursuant to 35 U.S.C. § 119(e), this application claims the benefit of U.S. Provisional Patent Application 62/590,150, filed Nov. 22, 2017, and U.S. Provisional Patent Application 62/596,446, filed Dec. 8, 2017, which are expressly incorporated herein by reference. This application is also related to U.S. patent application Ser. No. 15/______ (Attorney Docket No. KNL-001US01) and U.S. patent application Ser. No. 15/______ (Attorney Docket No. KNL-001US03), filed on the same date, which are expressly incorporated herein by reference.
- The present invention relates to methods and systems for non-invasive measurements in the human body, and in particular, methods and systems related to detecting physiologically dependent optical parameters in the human body.
- Measuring neural activity in the brain is useful for medical diagnostics, neuromodulation therapies, neuroengineering, or brain-computer interfacing. For example, it may be desirable to measure neural activity in the brain of a patient to determine if a particular region of the brain has been impacted by reduced blood irrigation, a hemorrhage, any other type of damage. For instance, in cases where the patient has suffered a traumatic brain injury, such as stroke, it may be desirable to determine whether the patient should undergo a therapeutic procedure. Measuring neural activity in the brain also may be used to determine the efficacy of such a therapeutic procedure.
- Conventional methods for measuring neural activity in the brain include diffuse optical tomography (DOT), and functional near-infrared spectroscopy (fNIRS), as well as others. These applications only employ a moderate amount of near-infrared or visible light radiation, thus being comparatively safe and gentle for a biological subject in comparison to X-Ray Computed Tomography (CT) scans, positron emission tomography (PET), or other methods that use higher-energy and potentially harmful radiation. Moreover, in contrast to methods, such as functional magnetic resonance imaging (fMRI), these optically-based imaging methods do not require large magnets or magnetic shielding, and thus, can be scaled to wearable or portable form factors, which is especially important in applications such as brain-computer interfacing.
- Because DOT and fNIRS rely on light, which scatters many times inside brain, skull, dura, pia, and skin tissues, the light paths occurring in these techniques comprise random or “diffusive” walks, and therefore, only limited spatial resolution can be obtained by a conventional optical detector, often on the order of centimeters. The reason for this limited spatial resolution is that the paths of photons striking the detector in such schemes are highly variable and difficult, and even impossible to predict without detailed microscopic knowledge of the scattering characteristics of the brain volume of interest, which is typically unavailable in practice (i.e., in the setting of non-invasive measurements through skull for brain imaging and brain interfacing). In summary, light scattering prevents optical imaging from achieving high resolution deep inside tissue.
- There is an increasing interest in ultrasound modulated optical tomography (UOT) to detect more precisely localized changes in biological tissues, e.g., on a sub-millimeter length scale, inside thick biological tissue, such as the brain (see U.S. Pat. No. 8,423,116; Sakadzic S, Wang L V, “High-Resolution Ultrasound-Modulated Optical Tomography in Biological Tissues,” Optics Letters, Vol. 29, No. 23, pp. 2770-2772, Dec. 1, 2004). These localized changes may include changes in light absorption in the brain that reflect neural activity and neurovascular coupling, such as a blood-oxygen-level dependent signal, for application in diagnostics, therapeutics, or, notably, brain computer interfacing (see Steinbrink J, Villringer A, Kempf F, Haux D. Boden S, Obrig H., “Illuminating the BOLD Signal: Combined fMRI-fNIRS Studies,” Magnetic Resonance Imaging, Vol. 24, No. 4, pp. 495-505, May 31, 2006). Thus, there is an increasing interest in ultrasound modulated optical tomography (UOT) in biomedical applications due to its potential to simultaneously achieve good resolution and imaging depth.
- In UOT, a highly localized ultrasound focus, e.g., millimeter or sub-millimeter in size, is used to selectively perturb (i.e., “tag”) light (e.g., light generated by a near-infrared coherent laser) passing through a voxel size of tissue defined by the size of the ultrasound focus. Due to the acousto-optic effect, light passing through the ultrasonic beam undergoes a frequency shift defined by multiples of the ultrasonic frequency. By detecting the frequency-shifted light, i.e., the tagged light, spatial information characterizing the biological tissue within the voxel can be acquired. As a result, spatial resolution is boosted from the centimeter-scale diffusive spread of light in the biological tissue to approximately a millimeter-scale voxel size. This ultrasound tagging of light relies on mechanisms known in the field (see Mahan G D, Engler W E, Tiemann J J, Uzgiris E, “Ultrasonic Tagging of Light: Theory,” Proceedings of the National Academy of Sciences, Vol. 95, No. 24, pp. 14015-14019, Nov. 24, 1998).
- Typical UOT implementations generate weak signals that make it difficult to differentiate ultrasound-tagged light passing through the focal voxel from a much larger amount of unmodulated light which is measured as DC shot noise. Thus, conventional UOT has the challenge of obtaining optical information through several centimeters of biological tissue, for example, noninvasive measurements through the human skull used to measure functional changes in the brain.
- Various methods have been developed to detect the very small fraction of tagged light out of a large background of untagged light by detecting the speckle pattern of light resulting from the interference of many multiply-scattered optical waves with different phases and amplitudes, which combine in a resultant wave whose amplitude, and therefore intensity, as well as phase, varies randomly. In the context of neuroengineering and brain computer interfacing, a key challenge is to render these methods to be sufficiently sensitive to be useful for through-human-skull functional neuroimaging.
- One technique uses a narrow spectral filter to separate out the untagged light striking a single-pixel detector, and is immune to speckle decorrelation (greater than ˜0.1 ms-1 ms) due to the scatters' motion (for example, blood flow) inside living biological tissue, but requires bulky and expensive equipment.
- Another technique uses crystal-based holography to combine a reference light beam and the sample light beam into a constructive interference pattern, but can be adversely affected by rapid speckle decorrelation, since the response time of the crystal is usually much longer than the speckle correlation time.
- Still another technique, referred to as heterodyne parallel speckle detection (PSD), employs optical interference together with a spatially resolved detector array (e.g., a conventional charge-coupled device (CCD) camera) used as an array of independent detectors for collecting the signal over a large number of coherence areas (see Atlan M, Forget B C, Ramaz F, Boccara A C, Gross M, “Pulsed Acousto-Optic Imaging in Dynamic Scattering Media With Heterodyne Parallel Speckle Detection,” Optics Letter, Vol. 30, No. 11, pp. 1360-1362, Jun. 1, 2005). Such configuration improves the signal-to-noise ratio relative to a single-detector and relative to approaches based on other modes of separating tagged and untagged light, such as spectral filters. However, the conventional CCD cameras used for heterodyne PSD have low frame rates, and therefore suffer from a relatively low speed relative to the speckle decorrelation time, thereby making this set up insufficient for in vivo deep tissue applications. Furthermore, conventional CCD cameras record both the AC signal and the DC background for each pixel. Thus, only a few bits of a pixel value can be used to represent the useful AC signal, while most of the bits are wasted in representing the DC background, resulting in a low efficiency in the use of bits.
- Lock-in cameras, as compared to conventional CCD cameras, have been used for comparatively bit-efficient and noise resistant heterodyne PSD (see Liu Y, Shen Y, Ma C, Shi J, Wang L V, “Lock-in Camera Based Heterodyne Holography for Ultrasound-Modulated Optical Tomography Inside Dynamic Scattering Media,” Applied Physics Letters, Vol. 108, No. 23, 231106, Jun. 6, 2016; see also Liu Y, Ma C, Shen Y, Wang L V, “Bit-Efficient, Sub-Millisecond Wavefront Measurement Using a Lock-In Camera for Time-Reversal Based Optical Focusing Inside Scattering Media,” Optics Letters, Vol. 41, No. 7, pp. 1321-1324, Apr. 1, 2016). For each pixel, a lock-in camera is capable of performing lock-in detection and outputting only information of the AC signal as a single AC amplitude map that is transferred to a computer, and thus, provides an efficient means of detecting and processing the speckle pattern.
- Besides the challenges posed by the signal-to-noise ratio, speckle decorrelation time, and efficient pixel bit processing, another challenge involves obtaining sufficient axial resolution (i.e., the depth resolution or ultrasound propagation direction). To address this challenge, UOT has been applied in a pulsed wave (PW) mode for heterodyne PSD, rather than a continuous (CW) mode (see Li Y Zhang H, Kim C, Wagner K H, Hemmer P., Wang L V, “Pulsed Ultrasound-Modulated Optical Tomography Using Spectral-Hole Burning as a Narrowband Spectral Filter,” Applied Physics Letters, Vol. 93, No. 1, 011111, Jul. 7, 2008; Ruan H, Mather M L, Morgan S P, “Pulsed Ultrasound Modulated Optical Tomography with Harmonic Lock-In Holography Detection,” JOSA A, Vol. 30, No. 7, pp. 1409-1416, Jul. 1, 2013).
- PW UOT has the benefit of enabling improved axial resolution compared to CW UOT. That is, with CW UOT, any light passing through the tissue, even though outside of the focal voxel, may be inadvertently tagged by the continuously propagating ultrasound energy along the ultrasound axis, thereby decreasing the signal-to-noise ratio. With PW UOT, the light passing through the tissue is pulsed only when the ultrasound pulses travels through the focal voxel, such that light outside of the focal voxel will not be tagged by the ultrasound energy. Although PW UOT improves axial resolution, the pulsed UOT signals are weak relative to continuous UOT signals.
- Although the UOT schemes described above may be sufficient for certain applications, such UOT schemes are inappropriate for the application of 3D-resolved, highly sensitive detection of small signals (e.g., blood-oxygen-level dependent signals) non-invasively through thick scattering layers, such as the human skull.
- In accordance with a first aspect of the present inventions, an ultrasound modulated optical tomography (UOT) system comprises an acoustic assembly configured for delivering ultrasound into a target voxel (e.g., one comprising brain matter) within an anatomical structure. The target voxel may be relatively small, e.g., less than one mm3.
- The UOT system further comprises an interferometer configured for delivering sample light into the anatomical structure, whereby a portion of the sample light passing through the target voxel is scattered by the anatomical structure as signal light, and another portion of the sample light not passing through the target voxel is scattered by the anatomical structure as background light that combines with the signal light to create a sample light pattern. The interferometer is further configured for combining reference light with the sample light pattern to generate an interference light pattern. The reference light may be combined with the signal light in a homodyne manner. For example, the interferometer may be further configured for frequency shifting the sample light by the frequency of the ultrasound, such that the reference light is combined with the signal light in the homodyne manner. In one embodiment, the interferometer comprises a light source configured for generating source light, a beam splitter configured for splitting the source light into the sample light and the reference light, and a light combiner configured for combining the reference light with the signal light and the background light to generate the interference light pattern.
- The UOT system further comprises a controller configured for operating the acoustic assembly and the interferometer to pulse the ultrasound and the sample light in synchrony, such that only the signal light is frequency shifted by the ultrasound. Multiple pulses of the sample light are delivered into the anatomical structure for each pulse of the ultrasound delivered into the target voxel. In one embodiment, the pulses of the sample light are identical. In this case, the interferometer may comprise at least one 1×N fiber splitter and at least one N×1 fiber coupler configured for generating the identical pulses of the sample light from a single optical pulse.
- The controller is further configured for operating the interferometer to sequentially modulate the interference light pattern to generate a plurality of different interference light patterns. In one embodiment, the interferometer is configured for sequentially modulating the interference light pattern by phase modulating the interference light pattern, e.g., by setting different phase differences (e.g., 0, π/2, π, and 3π/2) between sequential pulses of the sample light and the reference light. In this case, the interferometer may comprise an optical phase shifter configured for setting a phase difference between the sample light and the reference light to phase modulate the interference light pattern.
- In another embodiment, the controller is configured for operating the acoustic assembly and the interferometer to pulse the ultrasound and the sample light in synchrony, such that only a single pulse of the sample light is delivered into the anatomical structure for each pulse of the ultrasound delivered into the target voxel. In still another embodiment, the controller is configured for operating the acoustic assembly and the interferometer to pulse the ultrasound and the sample light in synchrony, such that multiple pulses of the sample light are delivered into the anatomical structure for each pulse of the ultrasound delivered into the target voxel.
- The UOT system further comprises at least one detector configured for detecting the different interference light patterns. In one embodiment, the detector(s) comprises an array of detectors configured for simultaneously detecting spatial components of each different interference light pattern. In this case, each detector respectively stores a plurality of values in a plurality of bins representative of the respective spatial components of the interference light patterns. Each of the interference light patterns may comprise a speckle light pattern, in which case, the spatial components may comprise speckle grains of the speckle light pattern. The array of detectors may be configured for simultaneously detecting spatial components of each different interference light pattern, and storing the plurality of values for all of the interference patterns in the plurality of bins within 10 milliseconds, and preferably within 1 microsecond to 1 millisecond. The UOT system may further comprise a lock-in camera that includes the array of detectors and corresponding bins.
- The UOT system further comprises a processor configured for determining a physiologically-dependent optical parameter (e.g., the level of deoxygenated and/or oxygenated hemoglobin concentration or relative abundance or the level of water concentration or relative water concentration of brain matter of brain matter) based on the plurality of values stored in the bins. In one embodiment, the processor is configured for determining neural activity within the target voxel based on the determined physiologically-dependent optical parameter. In one embodiment, the processor is configured for reconstructing the amplitude of the signal light using the plurality of values stored in each of the bins, and determining the physiologically-dependent optical parameter of the target voxel based on the reconstructed amplitude of the signal light. Each value may be respectively stored in each of the bins as an intensity of the spatial component of the respective interference light pattern, in which case, the processor may be configured for using the plurality of values stored in each of the bins to extract a product of the amplitude of the signal light and a known amplitude of the reference light, and determining the amplitude of the signal light from the extracted product.
- In accordance with a second aspect of the present inventions, a method of performing pulsed UOT comprises delivering ultrasound into a target voxel (e.g., one comprising brain matter) within an anatomical structure. The target voxel may be relatively small, e.g., less than one mm3.
- The method further comprises delivering sample light into the anatomical structure, whereby a portion of the sample light passing through the target voxel is scattered by the anatomical structure as signal light, and another portion of the sample light not passing through the target voxel is scattered by the anatomical structure as background light that combines with the signal light to create a sample light pattern.
- The method further comprises pulsing the ultrasound and the sample light in synchrony, such that only the signal light is frequency shifted by the ultrasound. Multiple pulses of the sample light are delivered into the anatomical structure for each pulse of the ultrasound delivered into the target voxel. In one method, the pulses of the sample light are identical. For example, the identical pulses of the sample light may be generated from a single optical pulse.
- The method further comprises combining reference light with the sample light pattern to generate an interference light pattern. The method may further comprise generating source light, and splitting the source light into the sample light and the reference light. The reference light may be combined with the signal light in a homodyne manner. For example, the method may further comprise frequency shifting the sample light by the frequency of the ultrasound, such that the reference light is combined with the signal light in the homodyne manner.
- The method further comprises sequentially modulating the interference light pattern to generate a plurality of different interference light patterns. In one method, the interference light pattern may be sequentially modulated by phase modulating the interference light pattern. For example, the interference light pattern may be phase modulated by setting different phase differences (e.g., 0, π/2, π, and 3π/2) between sequential pulses of the sample light and the reference light.
- The method further comprises detecting the different interference light patterns. In one method, spatial components of each different interference light pattern are simultaneously detected. In this case, the method may further comprise storing a plurality of values in a plurality of bins representative of the respective spatial components of the interference light patterns. Each of the interference light patterns may comprise a speckle light pattern, in which case, the spatial components may comprise speckle grains of the speckle light pattern. The spatial components of each different interference light pattern may be simultaneously detected, and the plurality of values for all of the interference patterns may be stored in the plurality of bins within 10 milliseconds, and preferably within 1 microsecond to 1 millisecond.
- The method further comprises determining a physiologically-dependent optical parameter (e.g., the level of deoxygenated and/or oxygenated hemoglobin concentration or relative abundance or the level of water concentration or relative water concentration of brain matter of brain matter) of the target voxel based on the plurality of values stored in the bins. One method further comprises determining neural activity within the target voxel based on the determined physiologically-dependent optical parameter.
- Other and further aspects and features of the invention will be evident from reading the following detailed description of the preferred embodiments, which are intended to illustrate, not limit, the invention.
- The drawings illustrate the design and utility of preferred embodiments of the present invention, in which similar elements are referred to by common reference numerals. In order to better appreciate how the above-recited and other advantages and objects of the present inventions are obtained, a more particular description of the present inventions briefly described above will be rendered by reference to specific embodiments thereof, which are illustrated in the accompanying drawings. Understanding that these drawings depict only typical embodiments of the invention and are not therefore to be considered limiting of its scope, the invention will be described and explained with additional specificity and detail through the use of the accompanying drawings in which:
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FIG. 1 is a block diagram of an ultrasound modulating optical tomography (UOT) system constructed in accordance with one embodiment of the present inventions; -
FIG. 2 a block diagram of one embodiment of an acoustic assembly used in the UOT system ofFIG. 1 ; -
FIG. 3a is a block diagram of one embodiment of an interferometer used in the UOT system ofFIG. 1 ; -
FIG. 3b is a block diagram of another embodiment of an interferometer used in the UOT system ofFIG. 1 ; -
FIG. 3c is a block diagram of still another embodiment of an interferometer used in the UOT system ofFIG. 1 ; -
FIG. 3d is a block diagram of yet another embodiment of an interferometer used in the UOT system ofFIG. 1 ; -
FIG. 4 is a schematic diagram of one embodiment of a detector array used in the UOT system ofFIG. 1 ; -
FIG. 5 is a timing diagram of one pulsing sequence used by the UOT system to detect a physiologically-dependent optical parameter in a target voxel within an anatomical structure; -
FIG. 6 is a schematic diagram of the UOT system ofFIG. 1 , particularly showing the generation of interference light patterns, the detection of spatial components in the, and binning of spatial component values; -
FIG. 7 is a timing diagram of another pulsing sequence used by the UOT system to detect a physiologically-dependent optical parameter in a target voxel within an anatomical structure; -
FIG. 8 is a plan diagram of one embodiment of an optical assembly used to split a single optical pulse into a train of identical optical pulses for use in the UOT system ofFIG. 1 ; -
FIG. 9a is a schematic diagram of a modified UOT system ofFIG. 1 , particularly showing a detection of a wavefront of signal light; -
FIG. 9b is a schematic diagram of a modified UOT system ofFIG. 1 , particularly showing playback of a phase conjugate of the wavefront of signal light; -
FIG. 10a is a block diagram of one embodiment of a phase conjugation array that can be incorporated into the UOT system ofFIG. 1 , particularly showing detection of the signal light; -
FIG. 10b is a block diagram of the phase conjugation array ofFIG. 10 a, particularly showing playback of a phase conjugation light field; -
FIG. 11 is a plot of the absorption of light in water over the wavelength of light; -
FIG. 12 is a plan view of wearable and unwearable units in which the UOT system ofFIG. 1 may be embodied; -
FIG. 13 is a flow diagram of one method used by the UOT system ofFIG. 1 to non-invasively measure a physiologically-dependent optical parameter using the pulse sequence ofFIG. 5 ; and -
FIG. 14 is a flow diagram of another method used by the UOT system ofFIG. 1 to non-invasively measure a physiologically-dependent optical parameter using the pulse sequence ofFIG. 7 . - The ultrasound modulated optical tomography (UOT) systems described herein utilize the combination of a pulsed ultrasound sequence that tags light propagating through an anatomical structure, and a selective lock-in camera that detects the tagged light (e.g., via parallel speckle detection (PSD)), as opposed to a conventional camera, to provide a highly efficient and scalable scheme that enables detection of highly localized and high spatial resolution UOT signals (e.g., blood-oxygen level dependent signals) at great depth inside a biological specimen, e.g., noninvasively through the entire thickness of the human skull and into the underlying cerebral cortical brain matter. The UOT systems may utilize a specific homodyne interference scheme that enables shot noise limited detection of the signal light. Such UOT signals may be used for, e.g., brain-computer interfacing, medical diagnostics, or medical therapeutics. Although the UOT systems are described herein as being used to image brain tissue for exemplary purposes, such UOT system can be used to image other anatomical parts of the body.
- Referring to
FIG. 1 , an ultrasound modulated optical tomography (UOT)system 10 constructed in accordance with one embodiment of the present inventions will be described. TheUOT system 10 is designed to non-invasively measure a physiologically-dependent optical parameter of atarget voxel 14 in ananatomical structure 16. In the illustrated embodiment, theanatomical structure 16 is the intact head of a patient 18 (shown inFIG. 12 ), including the scalp, skull, and brain, with thetarget voxel 14 being a portion of the brain. In a practical implementation, theUOT system 10 will acquire data from multiple target voxels 14 (“data voxels”) spatially separated from each other within a volume of interest (not shown). A “target voxel” may be defined as a small contiguous sub-volume of space (e.g., a cube) within theanatomical structure 16. For purposes of brevity, theUOT system 10 will be described as acquiring one data voxel (i.e., data representative of a physiologically-dependent optical parameter of the target voxel 14), although it should be understood that theUOT system 10 may be capable of acquiring more than one data voxel from the volume of interest of theanatomical structure 16. - In the illustrated embodiment, the physiologically-dependent optical parameter may be, e.g., a level of deoxygenated and/or oxygenated hemoglobin concentration or relative abundance, although in other embodiments, the physiologically-dependent optical parameter can be any parameter that varies in accordance with a change in an optical property of the target voxel 14 (e.g., light absorption). The physiologically-dependent optical parameters may alternatively comprise an analyte concentration in the blood, analyte/metabolite in tissue, concentration of a substance (e.g., blood, hemoglobin) or a structure within tissue, the presence and concentration of lamellar bodies in amniotic fluid for determining the level of lung maturity of a fetus, the presence and/or concentration of meconium in the amniotic fluid, optical properties of other extravascular fluids, such as pleural, pericardial, peritoneal, and synovial fluids. The physiologically-dependent optical parameter may be used internally within the
UOT system 10 or may be transmitted to external devices for use therein, e.g., medical devices, entertainment devices, neuromodulation stimulation devices, alarm systems, video games, etc. - The
UOT system 10 generally includes anacoustic assembly 20, aninterferometer 22, acontroller 24, a lock-incamera 28, and aprocessor 30. - The
acoustic assembly 20 is configured for deliveringultrasound 32 into thetarget voxel 14. Preferably, theacoustic assembly 20 focuses theultrasound 32 on thistarget voxel 14 in order to maximize the imaging resolution of theUOT system 10; that is, the more focused theultrasound 32 is, the smaller thetarget voxel 14 may be defined, thereby increasing the resolution of theUOT system 10. - Preferably, the frequency fus of the
ultrasound 32 is selected (e.g., in the range of 100 KHz-10 MHz), such that theultrasound 32 can pass efficiently through the skull and brain matter without significant attenuation that would otherwise cause insufficient ultrasound pressure at thetarget voxel 14, so that detectable UOT modulation of the light is created, as described in further detail below. It should be appreciated that the wavelength of such ultrasound in brain matter, given that the speed of sound in brain matter is similar to that of water (1500 meter/second), is on the order of fractions of a millimeter to a few millimeters. Thus, theacoustic assembly 20 may obtain ultrasound focal confinement at thetarget voxel 14 laterally on the order of the wavelength of the ultrasound 32 (e.g., less than 1 mm), and axially on the order of the wavelength of theultrasound 32 when theacoustic assembly 20 is operated to pulse theultrasound 32 at short durations (e.g., a single cycle), as will be described in further detail below. - Referring further to
FIG. 2 , one embodiment of theacoustic assembly 20 includes anultrasound transducer arrangement 34 and asignal generator 36. Theultrasound transducer arrangement 32 may take the form of any device that emits ultrasound 32 (in the illustrated embodiment, focused ultrasound) at a defined frequency and duration in response to a controlled drive signal; for example, signal acoustic element configured for emitting ultrasound beams with a fixed focus; or a piezoelectric phased array capable of emitting ultrasound beams with variable direction, focus, duration, and phase, or may be an array of pressure generating units (e.g., silicon, piezoelectric, polymer or other units), an ultrasound imaging probe, or even an array of laser generated ultrasound (LGU) elements. - The
signal generator 36 is configured for generating alternating current (AC) signals for driving theultrasound transducer arrangement 34 at a defined ultrasound frequency, duration, and intensity. The AC drive signal may be electrical or optical, depending on the nature of the ultrasound transducer arrangement. Thesignal generator 36 includes control inputs (not shown) for receiving control signals from thecontroller 24 that cause theultrasound transducer arrangement 34 to emit theultrasound 32 at a selected time, duration, and intensity. Thus, as will be described in further detail below, thecontroller 24 may selectively pulse theultrasound 32. - In one particular embodiment, the
transducer arrangement 34 is a head-mounted steerable ultrasonic array coupled to the skin of the patient via hydrogel or other means of mechanical coupling in order to effectively launch theultrasound 32 towards the precisely definedtarget voxel 14 within theanatomical structure 16, and in this case, the three-dimensional volume of the brain, while compensating the ultrasound wavefront using well-known phased array techniques to achieve efficient and selective ultrasound delivery to thetarget voxel 14. - Referring to
FIGS. 1 and 3 , theinterferometer 22 is configured for deliveringsample light 40 into theanatomical structure 16, where it scatters diffusively, e.g., through the human skull, into the brain, and back out again. Thus, aportion 40 a of thesample light 40 will pass through thetarget voxel 14 and will be scattered by theanatomical structure 16 assignal light 44, and anotherportion 40 b of thesample light 40 will not pass through thetarget voxel 14 and will be scattered by theanatomical structure 16 asbackground light 46. Thesignal light 44 andbackground light 44 combine together to create asample light pattern 47 that exits theanatomical structure 16. Theinterferometer 22 is further configured for combiningreference light 42 with thesample light pattern 47 to generate an interference light pattern 48 (e.g., speckle light pattern, which can be defined as an intensity pattern produced by the mutual interference of a set of scattered wavefronts; that is, a speckle light pattern results from the interference of many waves, but having different phases and amplitudes, which add together to give a result wave whose amplitude, and therefore intensity and phase, varies randomly). In the illustrated embodiment, theinterferometer 22 is configured for splitting source light 38 into thesample light 40 andreference light 42, as will be described in further detail below. - The
reference light 42 may be combined with thesignal light 44 in thesample light pattern 47 in a homodyne manner, e.g., by initially frequency shifting thesample light 40 by the frequency fus of theultrasound 32 delivered into thetarget voxel 14 by theacoustic assembly 20. That is, if unmodified, the samplelight portion 40 a passing through thetarget voxel 14 will be frequency shifted (i.e., tagged) by theultrasound 32 that also passes through thetarget voxel 14, such that thesignal light 44 will have frequencies f−fus. Presumably, the samplelight portion 40 b not passing through thetarget voxel 14 will not be frequency shifted (i.e., untagged) by theultrasound 32, such that thebackground light 46 will have a frequency f, i.e., the frequency of thesample light 40. It is also that not all of the samplelight portion 40 a passing through thetarget voxel 14 will be tagged by the ultrasound 32 (i.e., there exists a tagging efficiency (i.e., the number of tagged photons relative to a number of untagged photons scattered by the target voxel 14)), and therefore, some of the samplelight portion 40 a passing through thetarget voxel 14 will be scattered by theanatomical structure 16 asbackground light 46. - However, assuming that the
reference light 42 and thesample light 40 output by theinterferometer 22 have the same frequency f, in order to combine the ultrasound taggedsignal light 44 in thesample light pattern 47 and thereference light 42 in a homodyne manner, which requires thereference light 42 and signal light 44 to have the same frequency, the frequency f of thesample light 40 or thereference light 42 must initially be shifted relative to each other by the ultrasound frequency fus, such that, upon combining by theinterferometer 22, the frequency of the ultrasound taggedsignal light 44 will be shifted to the same frequency as thereference light 42, and the frequency of theuntagged background light 46 will differ from the frequency of thereference light 42 by the ultrasound frequency fus. Thus, either thesample light 40 or thereference light 42 will be pre-conditioned, such that the ultrasound taggedsignal light 44 will interfere with thereference light 42 in a homodyne manner, resulting in a DC interference component between thereference light 42 and signal light 44 that can be detected by the lock-incamera 28 as the signal component during each pulse, as will be described in further detail below. In contrast, the frequency shifting of thesample light 40 before it enters theanatomical structure 16, or the frequency shifting of thereference light 42, will prevent theuntagged background light 46 from interfering with thereference light 42 in a homodyne manner. - In the embodiment illustrated in
FIG. 1 , theinterferometer 22 down frequency shifts thesample light 40 by the ultrasound frequency fus to f−fus, such that the ultrasound taggedsignal light 44 has the frequency fus, theuntagged background light 46 has the frequency f−fus, and thereference light 42 has a frequency fus, thereby enabling combination of thereference light 42 and signal light 44 in a homodyne manner, as further described below with respect toFIG. 3 a. However, it is noted that because theultrasound 32 will tag thesignal light 44 with the ultrasound frequencies +fus and −fus, as well as other positive and negative multiples of the ultrasound frequency fus, other frequency shifting implementations are possible to effect homodyne combination of thereference light 42 and signal light 44. For example, as described in further detail below, theinterferometer 22 may up frequency shift thesample light 40 by the ultrasound frequency fus to f+fus, such that the ultrasound taggedsignal light 44 has the frequency fus, theuntagged background light 46 has the frequency f+fus, and thereference light 42 has a frequency fus (seeFIG. 3b ); may up frequency shift thereference light 42 by the ultrasound frequency fus to f+fus, such that the ultrasound taggedsignal light 44 has the frequency f+fus, theuntagged background light 46 has the frequency f, and thereference light 42 has a frequency f+fus (seeFIG. 3c ); may down frequency shift thereference light 42 by the ultrasound frequency fus to f−fus, such that the ultrasound taggedsignal light 44 has the frequency f−fus, theuntagged background light 46 has the frequency f, and thereference light 42 has a frequency f−fus (seeFIG. 3d ); or perform any other frequency shift of thesample light 40 orreference light 42 that results in the homodyne combination of thereference light 42 and thesignal light 44. - The
interferometer 22 is further configured for modulating (and in the illustrated embodiment, phase modulating) the interference light pattern to generate a plurality of different interference light patterns, which as will be described in further detail below, enables the amplitude of thesignal light 44 to be distinguished from thebackground light 46. - Referring further to
FIG. 3 a, one embodiment of theinterferometer 22 includes alight source 50, abeam splitter 52, anoptical phase shifter 54, anoptical frequency shifter 56, alight combiner 58, a pathlength adjustment mechanism 60, and a set ofmirrors - The
light source 50 is configured for generating coherent light as the source light 38, preferably at a single wavelength (e.g., in the range of 605 nm to 1300 nm), and may take the form of, e.g., a laser diode. In alternative embodiments, multiple light source(s) (not shown) may be used to generate the source light 38 at multiple distinct wavelengths, e.g., one generating source light 38 within the range of 605 nm to 800 nm, and another generating source light 38 within the range of 800 nm to 1300 nm. The coherence length of the source light 38 is preferably at least one meter in order to generate the best speckle contrast in thespeckle light pattern 48. Thelight source 50 may receive power from a drive circuit (not shown), which may include control inputs for receiving control signals from thecontroller 24 that cause thelight source 50 to emit the source light 38 at a selected time, duration, and intensity. Thus, as will be described in further detail below, thecontroller 24 may selectively pulse the source light 38, and thus thesample light 40 andreference light 42. - As specifically illustrated in
FIG. 3 a, thebeam splitter 52 is configured for splitting the source light 38 into thesample light 40 that propagates along a sample arm of theinterferometer 22 andreference light 42 that propagates along a reference arm of theinterferometer 22. In the illustrated embodiment, the beam splitter 52 (e.g., a partially transparent mirror) splits the source light 38 via amplitude division by reflecting a portion of the source light 38 as thesample light 40, and transmitting the remaining portion of the source light 38 as thereference light 42, although thebeam splitter 52 may alternatively reflect a portion of the source light 38 as thereference light 42, and transmit the remaining portion of the source light 38 as thesample light 40. In alternative embodiments, thebeam splitter 52 may split the source light 38 via wavefront division by splitting a portion of the wavefront into thesample light 40 and splitting the remaining portion of the wavefront into thereference light 42. In either case, thebeam splitter 52 may not necessarily split the source light 38 equally into thesample light 40 andreference light 42, and it may actually be more beneficial for thebeam splitter 52 to split the source light 38 unevenly, such that the amplitude of thesample light 40 is less than the amplitude of the reference light 42 (e.g., 10/90 power ratio) in order to comply with tissue safety standards. That is, the amplitude of thesample light 40 will preferably be relatively low to avoid damaging the tissue, whereas the amplitude of thereference light 42, which will be used to boost thesignal light 44 in theinterference light pattern 46, will be relatively high. - The
optical phase shifter 54 is configured for setting the phase difference between thesample light 40 andreference light 42. Theoptical phase shifter 54 may include control inputs (not shown) for receiving control signals from thecontroller 24 that cause theoptical phase shifter 54 to set the phase of thereference light 42 relative to thesample light 40. Thus, as will be described in further detail below, thecontroller 24 may selectively set the phase between thesample light 40 and thereference light 42. - The
optical frequency shifter 56 is configured for down frequency shifting thesample light 40 by the ultrasound frequency fus to f−fus, such that the frequency of the ultrasound taggedsignal light 44 will be f, while the frequency of thebackground light 46 will be f−fus, thereby enabling the homodyne combination of thereference light 42 at frequency f and the ultrasound taggedsignal light 44 at frequency f, as described above with respect toFIG. 1 . In one alternative embodiment illustrated inFIG. 3 b, theoptical frequency shifter 56 is configured for up frequency shifting thesample light 40 by the ultrasound frequency fus to f+fus, such that the frequency of the ultrasound taggedsignal light 44 will be f, while the frequency of thebackground light 46 will be f +fus, thereby enabling the homodyne combination of thereference light 42 at frequency f and the ultrasound taggedsignal light 44 at frequency f. In one alternative embodiment illustrated inFIG. 3 c, theoptical frequency shifter 56 is configured for up frequency shifting thereference light 42 by the ultrasound frequency fus to f+fus, such that the frequency of the ultrasound taggedsignal light 44 will be f+fus, while the frequency of thebackground light 46 will be f, thereby enabling the homodyne combination of thereference light 42 at frequency f+fus and the ultrasound taggedsignal light 44 at frequency f+fus. In still another alternative embodiment illustrated inFIG. 3 d, theoptical frequency shifter 56 is configured for down frequency shifting thereference light 42 by the ultrasound frequency fus to f−fus, such that the frequency of the ultrasound taggedsignal light 44 will be f−fus, while the frequency of thebackground light 46 will be f, thereby enabling the homodyne combination of thereference light 42 at frequency f−fus and the ultrasound taggedsignal light 44 at frequency f−fus. - In any event, the
frequency shifter 54 may include a local oscillator (not shown) that outputs a signal having a fixed or variable frequency. The local oscillator may be variable, in which case, it may have a control input for receiving control signals from thecontroller 24 that cause the local oscillator to output a signal at a defined frequency. Alternatively, the local oscillator may be fixed, in which case, it will output a signal having a fixed frequency. In either case, the frequency of the signal output by the local oscillator will be equal to the frequency fus of theultrasound 32 emitted by theacoustic assembly 20. - The
light combiner 58 is configured for combining thereference light 42 with thesample light pattern 47 via superposition to generate theinterference light pattern 48. Thelight combiner 58 can take the form of, e.g., a combiner/splitter mirror. - The path
length adjustment mechanism 60 is configured for adjusting the optical path length of the reference light 42 (i.e., the reference arm) to nominally match the expected optical path length of the combinedsample light 40 and signal light 44 (i.e., the sample arm), such that thesignal light 44 and thereference light 42 reach thelight combiner 58 at the same time. The pathlength adjustment mechanism 60 may include a beam splitter/combiner 64 and anadjustable mirror 66 that can be displaced relative to the beam splitter/combiner 64. The beam/splitter combiner 64 is configured for redirecting thereference light 42 at a ninety-degree angle towards themirror 66, and redirecting thereference light 42 reflected back from themirror 66 at a ninety-degree angle towards thelight combiner 58. Thus, adjusting the distance between themirror 66 and the beam splitter/combiner 64 will adjust the optical path length of the reference arm to match the optical path length of the sample arm. - The mirror assembly 62 is configured for confining the optical light paths in the
interferometer 22 into a small form factor, and in the illustrated embodiment, includes a first tilted, completely reflective, mirror 62 a configured for redirecting thesample light 40 at a ninety-degree angle towards thebiological specimen 16, and a second tilted, completely reflective,mirror 62 b configured for redirecting the signal light 44 (and coincidentally a portion of the background light 46) towards thelight combiner 58. - Referring back to
FIG. 1 , thecontroller 24, which may, e.g., take the form of a central processing unit (CPU), is configured for implementing pulsed wave (PW) UOT by operating theacoustic assembly 20 to pulse the ultrasound 32 (in the illustrated embodiment, by sending on/off control signals to the signal generator 36), and operating theinterferometer 22 to pulse the sample light 40 (in the illustrated embodiment, by sending on/off control signals to the drive circuit coupled to the light source 50) in synchrony with the (comparatively very slow) flight of theultrasound 32, such that only thesignal light 44 is frequency shifted (i.e., tagged) by theultrasound 32. That is, a pulse of thesample light 40 will be delivered into theanatomical structure 16, such that it will pass through thetarget voxel 14 only as theultrasound 32 passes through thetarget voxel 14. In this manner, no portion of thebackground light 46 will be tagged by theultrasound 32. As a result, pulsed wave (PW) UOT improves the spatial resolution in the axial direction (or depth) compared to continuous wave (CW) UOT. Thus, PW UOT achieves axial confinement and three-dimensional (3D) spatial resolution, rather than merely two-dimensional (2D) spatial resolution as in the case with CW UOT. - The
controller 24 is further configured for operating theinterferometer 22 to sequentially modulate the interference light pattern 48 (in the illustrated embodiment, by sending on/off control signals to the optical phase shifter 54) to generate a plurality of different interference light patterns. As will be described in further detail below, theinterferometer 22 will set different phases (and in the illustrated embodiment, four different phases equal to 0, π/2, π, and 3π/2) between sequential pulses of thesample light 40 and thereference light 42 to facilitate quadrature detection of thesignal light 44. - As will be also described in further detail below, the
controller 24 is further configured for synchronously operating the lock-incamera 28, such that the bin shifting of data detected by the lock-incamera 28 is performed in synchrony with the phase changes in theinterferometer 22. - Referring further to
FIG. 4 , the lock-incamera 28 includes an array of detectors 68 (or “pixels”) configured for simultaneously detecting spatial components of each of the different interferencelight patterns 48. In the case where theinterference light pattern 48 is a speckle light pattern, the spatial components are speckle grains (approximately the size of a wavelength of the light) of the speckle light pattern. In general, lock-in cameras include a class of digital cameras in which multiple measurements of a light field are rapidly made at each pixel in a temporally precise fashion synchronized with an external trigger or oscillation and stored in multiple “bins” within each pixel, in contrast with conventional cameras, which store only one value per pixel that merely aggregate the incoming photo-electrons over the camera frame integration time. Lock-in cameras may also perform on-chip computations on the binned values. Thus, the key feature of lock-in cameras is their ability to rapidly capture and store multiple sequential samples of the light field, with sample-to-sample latencies shorter than readout times of conventional cameras. This feature enables them, for example, to sample a modulated light field at the same frequency as the modulation, such that subtraction across successive samples, or other operations, such as quadrature detection (discussed below) will extract the component of the light that is modulated at the modulation frequency, while subtracting off the unmodulated (“DC”) background. Similarly, lock-in cameras can be used to make a series of such measurements or comparisons, locked to an external trigger signal (generated by the controller 24), rapidly in order to extract such modulated components from a rapidly changing light field arising from a dynamic, disordered biological specimen. - Thus, each
detector 68 of the lock-incamera 28 respectively stores a plurality of values in a plurality of bins 70 a-70 d representative of the spatial component of the fourinterference light patterns 48, and in this case, four bins 70 a-d (in general, 70) for storing four values from the respective fourinterference light patterns 48. The spatial component values stored in the bins 70 of arespective detector 68 may be, e.g., the intensity values of the respective spatial component ofinterference light patterns 48. For example, for any particular detector 68 (or pixel) corresponding to a particular spatial component (or speckle grain), four power values Pa-Pd for the fourinterference patterns 48 will be respectively stored in the four bins 70 a-70 d. As will be described in further detail below, the spatial component power values Pa-Pd detected by eachdetector 68 of thecamera 28 for the fourinterference patterns 48 can be used to reconstruct the amplitude of thesignal light 44, and thus, can be said to be representative of the physiologically-dependent optical parameters (e.g., optical absorption) of thetarget voxel 14. The lock-incamera 28 includes control inputs (not shown) for receiving control signals from thecontroller 24, such that the detection and binning of the data can be coordinated with the pulsing of theultrasound 32 and sample light 40 described in further detail below. - Although only a single lock-in
camera 28 is illustrated, it should be appreciated that multiple lock-in cameras 28 (e.g., in an array) or a lock-in camera in the form of multiple camera sensor chips on a common circuit board, can be used to increase the number of detectors 68 (i.e., pixels). Although not illustrated, thesystem 10 may include magnification optics and/or apertures to magnify the individual speckle grains, which may have a size on the order of the wavelength of the near-infrared or visible light used to acquire the data voxel, and hence on the order of hundreds of nanometers in size, to approximately the sizes of thedetectors 68 of the lock-incamera 28. Thus, in the illustrated embodiment, the pixel sizes and pitches of the lock-incamera 28 are matched to the speckle grain sizes and pitches of theinterference light pattern 48 via the appropriate magnification, although other embodiments are possible. - Referring to
FIGS. 5 and 6 , one pulsing sequence that can be used in a PW UOT technique performed by thesystem 10 for generating fourinterference light patterns 48 and detecting and storing the spatial component power values for theinterference light patterns 48 will be described. - During one acquisition of a single data voxel (i.e., acquisition of data characterizing the target voxel 14), an ultrasound pulse train consisting of four separate, but identical, ultrasound pulses Ua-Ud are delivered into the
target voxel 14. In this embodiment, the duration τ of each ultrasound pulse U is equal to only one full cycle of theultrasound 32 to maximize the data acquisition speed, and thus, is equal to 1/fus, although in alternative embodiments, the duration τ may be several ultrasound cycles long (e.g., on the order of 1 microsecond or less than one microsecond). It should be noted that it is desirable to minimize the duration τ of the ultrasound pulse U in order to minimize ultrasound focal confinement at thetarget voxel 14. - The duty cycle of the ultrasound pulses Ua-Ud (i.e., the time that elapses between the beginning of one pulse U to the beginning of the next pulse U) is τduty. The duty cycle τduty may be selected to allow each ultrasound pulse U to exit the
anatomical structure 16 before the next measurement is taken, such that the ultrasound taggedsignal light 44 is only present at high pressures at the three-dimensional location of thetarget voxel 14. The frame rate of the lock-incamera 28 should be selected to match the duty cycle τduty of the ultrasound pulse U, such that there exists one ultrasound pulse U per frame. - A light pulse train consisting of four sample light pulses La-Ld is also delivered into the
anatomical structure 16 in synchrony with the delivery of the four ultrasound pulses Ua-Ud, such that, as each ultrasound pulse U passes through thetarget voxel 14, the sample light pulse L likewise passes through thetarget voxel 14. - In this manner, only the signal light 44 (and none of the background light 46) is tagged with the ultrasound, as discussed above. In this particular embodiment, only one sample light pulse L is delivered for each ultrasound pulse U. Thus, there is a one-to-one correspondence between the sample light pulses La-Ld and the ultrasound pulses Ua-Ud. Although each of the sample light pulses L is illustrated in
FIG. 5 as having the same width as the width of the respective ultrasound pulses U for purposes of illustration, each sample light pulses L is, in practicality, at least slightly smaller than the respective ultrasound pulse U due to the latency period required for the respective ultrasound pulse to reach thetarget voxel 14. Alternatively, the duration of the sample light pulse L, the duration of the sample light pulse L can be much less than the duration of the ultrasound pulse U. In any event, the duration of the sample light pulse L preferably should be approximately matched to be an integer multiple of the frequency of the ultrasound frequency fus, e.g., if the ultrasound frequency fus of 2 MHz, the duration of the sample light pulse L should be multiples of 0.5 microseconds. AlthoughFIG. 5 illustrates only one ultrasound cycle per ultrasound pulse U, if the duration of the sample light pulse L is multiple integers of the frequency of the ultrasound frequency fus, the number ultrasound cycles per ultrasound pulse U is preferably equal to the same multiple integer. This ensures that a full, balanced cycle of tagged light is generated. The energy of each sample light pulse L should be sufficiently high, e.g., on the order of 1 microsecond in duration (but can be as low as 10 nanoseconds in duration) and tens of micro-Joules per square centimeter. - For each of the four separate ultrasound pulses Ua-Ud occurring during the acquisition of a single data voxel, the phase difference between the
reference light 42 and thesample light 40 is set to a different setting, and in this case, to one of 0, π/2, π, and 3π/2. In the illustrated embodiment, the phase between thereference light 42 and thesample light 40 is sequentially set to 0, π/2, π, and 3π/2, although these phase settings can be performed in any order, as long as all fourphase settings 0, π/2, π, and 3π/2 are used during the acquisition of a single data voxel. - The respective pulses of the
sample light pattern 47 andreference light 42 are then combined into theinterference light patterns 48, each having four corresponding interference pulses Ia-Id that can be detected by the lock-incamera 28. That is, for each interference pulse I, adetector 68 detects a spatial component of the respective interference pulse I (e.g., a speckle grain in the case where theinterference light pattern 48 includes a speckle pattern) and stores the spatial component value (e.g., power) within a respective one of the bins 70. - That is, at phase φ=0, a given pixel n will detect and store the value of the respective spatial component of the interference pulse Ia into
bin 1 of the pixel n; at phase φ=π/2, the pixel n will detect and store the value of the respective spatial component of the interference pulse Ib intobin 2 of the pixel n; at phase φ=π, the pixel n will detect and store the value of the respective spatial component of the interference pulse Ic intobin 3 of the pixel n; and at phase φ=3π/2, the pixel n will detect and store the value of the respective spatial component of the interference pulse Id intobin 4 of the pixel n. - Similarly, at phase φ=0, the next pixel n+1 will detect and store the value of the respective spatial component of the interference pulse Ia into
bin 1 of the pixel n+1; at phase φ=π/2, the pixel n+1 will detect and store the value of the respective spatial component of the interference pulse Ib intobin 2 of the pixel n+1; at phase φ=π, the pixel n+1 will detect and store the value of the respective spatial component of the interference pulse intobin 3 of the pixel n+1; and at phase φ=3π/2, the pixel n+1 will detect and store the value of the respective spatial component of the interference pulse Id intobin 4 of thepixel n+ 1. - Similarly, at phase φ=0, the next pixel n+2 will detect and store the value of the respective spatial component of the interference pulse Ia into
bin 1 of the pixel n+2; at phase φ=π/2, the pixel n+2 will detect and store the value of the respective spatial component of the interference pulse Ib intobin 2 of the pixel n+2; at phase φ=π, the pixel n+2 will detect and store the value of the respective spatial component of the interference pulse intobin 3 of the pixel n+2; and at phase φ=3π/2, the pixel n+2 will detect and store the value of the respective spatial component of the interference pulse Id intobin 4 of thepixel n+ 2. - Thus, for each of an n number of pixels, four values will be respectively stored in the four bins 1-4. Significantly, in the case where the
interference light pattern 48 includes a speckle light pattern, it is important that all four sample light pulses P be delivered by theinterferometer 22 to thetarget voxel 14 and that all four interference pulses I be detected and recorded by thecamera 28 within the characteristic speckle decorrelation time of thetarget voxel 14, which scales super-linearly with the depth into theanatomical structure 16 at which thetarget voxel 14 is located. For imaging deep inside a living biological tissue, such as through the human skull and into the human cerebral cortex, the speckle decorrelation time is expected to be on the order of microseconds to tens of microseconds. For imaging directly into living brain matter in the absence of skull, speckle decorrelation times have been measured to be on the order of ten milliseconds for 1-millimeter penetration or 1-millisecond for 3-millimeter penetration. Notably, the speckle decorrelation time impacts the depth scaling of lock-in camera based UOT in dynamic scattering media, such as biological tissue, namely the constraint that multiple phase-shifted measurements must be made within the speckle decorrelation time (see, e.g., Qureshi M M, Brake J., Jeon H J, Ruan H, Liu Y, Safi A M, Eom T J, Yang C., Chung E, “In Vivo Study of Optical Speckle Decorrelation Time Across Depths in the Mouse Brain,” Biomedical Optics Express, Vol. 8, No. 11, pp. 4855-4864 (Nov. 1, 2017). Thus, it is important that the time window in which the set of quadrature measurements is short enough that thetarget voxel 14 does not have the time to de-correlate significantly. Otherwise, the signal-to-noise ratio is diminished. - Referring to
FIG. 7 , another particularly advantageous pulsing sequence that can be used in a PW UOT technique performed by thesystem 10 for generating fourinterference light patterns 48 and detecting and storing the spatial component power values for theinterference light patterns 48 will be described. The pulsing sequence ofFIG. 7 is identical to the pulsing sequence ofFIG. 5 , with the exception that multiple sample light pulses L, and in this embodiment, all four sample light pulses L, are delivered to thetarget voxel 14 for each ultrasound pulse U delivered to thetarget voxel 14, thereby accelerating the delivery of the sample light pulses P by theinterferometer 22 to thetarget voxel 14, and the resultant generation, detection, and recording of all four interference pulses I by thecamera 28. That is, because the four sample light pulses L are delivered for each ultrasound pulse U, the speed of the data voxel acquisition is increased by a factor of four. - In particular, during the acquisition of four consecutive data voxels (as opposed to only one in the pulsing sequence of
FIG. 5 ), an ultrasound pulse train consisting of four separate, but identical, ultrasound pulses U are delivered into thetarget voxel 14. As with the case in the pulsing sequence ofFIG. 5 , the duration τ of this ultrasound pulse U is equal to only one full cycle of theultrasound 32 to maximize the data acquisition speed, and thus, is equal to 1/fus. The duration τ of this ultrasound pulse U and the duty cycle τduty of the ultrasound train pulse ofFIG. 7 can be identical to the respective duration τ and duty cycle τduty of the ultrasound pulse train in the pulsing sequence ofFIG. 7 . - A light pulse train consisting of four sets of sample light pulses, with each set comprising four sample light pulses La-Ld, are also delivered into the
anatomical structure 16 in synchrony with the delivery of the four ultrasound pulses U, such that as each ultrasound pulse U passes through thetarget voxel 14, the corresponding set of four sample light pulses La-Ld, likewise pass through thetarget voxel 14. Thus, only the signal light 44 (and none of the background light 46) is tagged with the ultrasound, as discussed above. Thus, four sample light pulses La-Ld are delivered for each ultrasound pulse U. Thus, there is a four-to-one correspondence between the sample light pulses La-Ld and ultrasound pulses U. - Thus, in the same manner described above with respect to the pulsing sequence illustrated in
FIG. 5 , for each ultrasound pulse U occurring during the acquisition of a single data voxel, the phase difference between thereference light 42 and thesample light 40 is set to a different setting, and in this case, to one of 0, π/2, π, and 3π/2, to generate four interference pulses Ia-Id. The quick detection and storage scheme of the lock-incamera 28 enables acquisition of an entire data voxel within one cycle of theultrasound 32, well within the speckle decorrelation time of thetarget voxel 14. - It can be appreciated that the use of the lock-in
camera 28 provides for a high-speed and precisely timed detection method that can capture differences in a light field far faster than the frame rates of conventional cameras. In the illustrated embodiment, the lock-incamera 28 rapidly measures the four quadratures of the pulse sequences illustrated inFIGS. 5 and 7 . The acquisition sequence can be precisely timed to external signals for integration with optimal ultrasound and light pulse sequences. The lock-incamera 28 also enables an efficient detection scheme compared to conventional pulsed UOT, since the pulsed UOT signals may be quickly detected with a high signal-to-noise ratio, while using the full bit depth of the analog-digital conversion available in the signal chain due to rejection of DC background by the lock-incamera 28. The lock-incamera 28 provides for a simplified detection scheme that is highly scalable to large numbers of pixels and high frame rates, enabling the maximization of signal capture in UOT, thereby improving spatial and temporal resolution. - It should be appreciated that in addition to the ability of the combination of the pulsed UOT with a lock-in camera to provide high axial spatial resolution and high sensitivity from the high-speed lock-in detection, such combination also provides the additional advantage of efficiently detecting the signal light associated with a specific time point on the ultrasound phase cycle (e.g., at the peaks of the ultrasound phase cycle). As such, the pulsed UOT/lock-in camera combination can accurately image tissue with a relatively small number of data measurements, and thus, a relatively short period of time, preferably within the speckle decorrelation time of the target voxel. In comparison, a continuous wave approach results in averaging light signal detection over a range of arbitrarily placed points on the ultrasound phase cycle, leading to a diminished overall detection sensitivity, requiring that, for sufficient sensitivity, data measurements be taken over a period time longer than the speckle decorrelation time of the target voxel. Thus, the use of pulsed UOT in combination with the lock-in camera allows deeper imaging into tissue.
- The detection processes illustrated in
FIGS. 5 and 7 require the ultrasound timing and intensity to be consistent. As such, the output of theacoustic assembly 20 may be periodically sampled to ensure that thesystem 10 does not suffer from ultrasound performance drift, for instance, as thetransducer arrangement 34 heats. Furthermore, the detection processes illustrated inFIGS. 5 and 7 require that all of the light pulses used between phase changes in the quadrature measurements be equal in strength, or at least exhibit a known ratio that can be normalized out post-detection. In the case that the pulse-to-pulse variations in the light intensity emitted by thelight source 50 are too large, the optical arrangement illustrated inFIG. 8 can be utilized to generate optical pulses that are identical in intensity and intensity-time profile, but temporally separated from each other. - In particular, this optical arrangement includes a first 1×2
fiber splitter 72 a in which a single optical pulse P (generated by the light source 50) is input via anoptical fiber 74 and split into two identical optical pulses P1, P2. Twooptical fibers fiber splitter 72 a, such that the two identical optical pulses P1, P2 respectively propagate within the twooptical fibers fiber coupler 76 a into which the two identical optical pulses P1, P2 are input and combined, and output to a singleoptical fiber 74 c. By making the lengths of theoptical fibers fiber splitter 72 a is effectively split into two identical optical pulses that propagate through the singleoptical fiber 74 c and are spaced out by a time difference determined by the optical path length difference between the twooptical fibers - Another fiber coupler and pair of optical fibers can be added to create four identical optical pulses separated from each other in time. In particular, the optical arrangement further includes a second 1×2
fiber splitter 72 b to which the singleoptical fiber 74 c carrying the two identical and temporally spaced optical pulses P1, P2 is coupled. Thus, the two identical optical pulses P1, P2 are input into the second 1×2fiber splitter 72 b and split into four identical optical pulses P1 a, P1 b, P2 a, P2 b (i.e., the optical pulse P1 is split into optical pulses P1 a, P1 b, and the optical pulse P2 is split into optical pulses P2 a, P2 b). Twooptical fibers fiber splitter 72 b, such that the two sets of two identical optical pulses P1 a, P1 b and P2 a, P2 b respectively propagate within the two optical fibers 72 d, 72 e. The optical arrangement further includes a second 2×1fiber coupler 76 b into which the two sets of identical optical pulses P1 a, P1 b and P2 a, P2 b are input and combined, and output to a singleoptical fiber 74 f. By making the lengths of theoptical fibers fiber splitter 72 b are effectively split into four identical optical pulses that propagate through the singleoptical fiber 74 f and spaced out by a time difference determined by the optical path length difference between the twooptical fibers - Referring back to
FIG. 1 , once thecamera 28 acquires the data voxel by storing all spatial component values of each of the four interference pulses Ia-Id within the four bins 70 of each of thedetectors 68, these data can be sent to the processor 30 (which can, e.g., take the form of a computer, field-programmable gate array or application specific integrated circuit), which is configured for determining a physiologically-dependent optical parameter (e.g., absorption) of thetarget voxel 14 based on the four values stored in the bins 70 of eachdetector 68. As briefly discussed above, the four spatial component values can be power values Pa-Pd, which can be used by theprocessor 30 to reconstruct the amplitude of thesignal light 44, and thus, can be said to be representative of the physiologically-dependent optical parameters (e.g., optical absorption) of thetarget voxel 14. - The spatial component power values Pa-Pd for all four interference light patterns Ia-Id can be used in accordance with known “quadrature detection” methods to reconstruct the amplitude of the
signal light 44, which is proportional to the number of tagged photons emerging from the target voxel 14 (i.e., the number of photons in the signal light 44), and thus, can be used to measure optical absorption in the target voxel 14 (e.g., for the purpose of measuring spatially localized neural activity-correlated changes in the level of deoxygenated and/or oxygenated hemoglobin concentration or relative abundance in the brain, which appear as localized changes in the optical absorption of blood). In the illustrated embodiment, it should be understood that because of the diffusive scattering of light over large distances through the brain and skull, theinterference light pattern 48 detected by the lock-incamera 28 takes the form of a random speckle pattern in which each localized speckle grain has a definite, but random phase offset in the interference light pattern 48 (i.e., a beat pattern) between thereference light 42 and thesignal light 44. This results in the unknown random phases in the beat patterns measured by each detector 68 (or pixel) in the equations set forth below. - In particular, the power detected at a single detector 68 (or pixel) for each optical pulse at one of the four phases can be expressed as:
- [1] Value1,k=Pbackground+Psignal+Preference+2(Psignal×Preference)1/2×cos(φcontrol−φunknown1, speckle k)+2(Psignal×Pbackground)1/2×cos(2π×fus−φunknown2, speckle k)+2(Preference×Pbackground)1/2×cos(2π×fus−φunknown3, speckle k), where Pbackground represents light at frequency f−fus that has not been tagged with the
ultrasound 32; Psignal represents light at frequency f that has been tagged with theultrasound 32; Preference represents the reference light at frequency f; φcontrol is a control phase shift introduced into thereference light 42 for each detectedinterference pattern 48; φunknown1, speckle k, φunknown2, speckle k, and φunknown3, speckle k are random phases at the kth speckle grain at the time of measurement, which originates via multiple scattering of coherent light inside the tissue. - The terms Pbackground+Psignal+Preference are constant across all four optical pulses with different control phase values φcontrol. The terms 2(Psignal×Pbackground)1/2×cos(2π×fus−φunknown2)+2(Preference×Pbackground)1/2×cos(2π×fus−φunknown3) oscillate at the frequency fus, and are not detected by the lock-in
camera 28, and thus, can be ignored. As such, equation [1] can be reduced to: - [2] Pbackground+Psignal+Preference+2(Psignal×Preference)1/2×cos(φcontrol−φunknown), which is analogous to the well-known quadrature formula:
- [3] A+B×Cos(φcontrol+φunknown), where φcontrol can be respectively set to 0, π/2, π, and 3π/2 to create four equations. Both the amplitude B and the unknown phase φunknown can be extracted by solving the resulting four equations using the standard trigonometric identities.
- Thus, the term magnitude of Psignal×Preference can be extracted by shifting the control phase φcontrol successively on each of four successive pulses φcontrol=0, π/2, π, and 3π/2. Even though φunknown is an unknown and random phase, specific to each pixel, which results from the laser speckle pattern due to light scattering in the tissue, by measuring and storing each of these four measurements at different control phase values φcontrol, the value of the interference term 2(Psignal×Preference)1/2 may be extracted via the known principal of “quadrature detection.” Because the power of the reference signal Preference is known or independently measurable, the interference term 2(Psignal×Preference)1/2 serves as a measurement of the power of the signal light Psignal. Thus, using a known scaling relationship, the power of the signal light Psignal can be determined (either in the absolute sense or relative sense) from the extracted term interference term 2(Psignal×Preference)1/2.
- Because the speckle phases are random, according to the known principles of parallel speckle detection in UOT or in wavefront measurement from strongly scattering media, it is known that a single-pixel detector will not scale to high signal to noise ratios. In particular, the aggregate signal over a large single-pixel detector would scale as the square root of detector size, but so would shot noise in the background, and hence the signal to noise ratio performance of a large detector would not increase with detector size. In contrast, as described in the equations below, with lock-in detection at each detector (or pixel), the aggregate signal scales linearly with the number of pixels, while the aggregate background shot noise scales as the square root, and hence signal to noise performance increases as the square root of the number of pixels, giving a strong advantage for using large numbers of pixels.
- It can be assumed that the amplitude of Preference is much greater than the amplitude of Pbackground, and the amplitude of Psignal is naturally much less than the amplitude of Preference, since the ultrasound tagged
signal light 44 originates from a verysmall target voxel 14 within the tissue and the tagging efficiency (i.e., the number of tagged photons relative to a number of untagged photons scattered by the target voxel 14) within thattarget voxel 14 is a small fraction. Thus, only interference terms containing Preference are significant in the sum representing the intensity measured by each pixel (i.e., Pbackground+Psignal+Preference+2(Psignal×Preference)1/2×cos(φcontrol−φunknown1)). - Therefore, the dominant signal source contributing to detection has the following number of photons impinging on one pixel:
- [4] dominant signal=(ε/hv)×2(Psignal×Preference)1/2τ;
and the dominant noise source in the quadrature measurement of this amplitude is due to the shot noise in thereference light 42, and has the following number of photons impinging on each pixel: - [5] dominant noise=((ε/hv)×Preference×τ)1/2;
where ε is a detector efficiency scaling factor, P is the power for each of the ultrasound tagged photons, hv is the per-photon energy (with h as Plank's constant, and v as the frequency of the light), and τ is the integrated pulse widths used in the measurement. - With a number of pixels N, the signal-to-noise ratio (SNR) scales with N1/2, since the total shot noise grows as N1/2, whereas the total signal grows as N, so that:
- [6] SNRN pixels=(N×(ε/hv)×τ×Psignal)1/2, which shows that the SNR improves with increasing number of pixels in the lock-in
camera 28. Thus, the Poisson nature of photon shot noise statistics is being utilized to determine the fundamental signal to noise ratio. - It should be appreciated that although the
UOT system 10 has been described as using a 4-bin quadrature detection scheme to reconstruct the amplitude of the signal light 44 from theinterference light patterns 48, and therefore, utilizes four bins 70 (and four optical pulses) for each detector 68 (or pixel) of the lock-incamera 28 to store the intensity values of the respective fourinterference patterns 48 over the four different phases, theUOT system 10 may utilize less than four phases (e.g., three phases equal to 0, 2π/3, and 4π/3), or may even utilize two phases (e.g., 0 and π) to reconstruct the amplitude of the signal light 44 from theinterference light patterns 48, and therefore utilizes three bins 70 (and three optical pulses) or only two bins 70 (and only two optical pulses) for each detector 68 (or pixel) to store the intensity values of therespective interference patterns 48 over the phases. It should further be appreciated that although the phases of the 4-bin quadrature scheme, as well as the three-bin and two-bin detection schemes, have been described as being equally spaced, the phases used in any of these detection schemes can be unequally spaced. For example, for the three-bin detection scheme, the phases can be selected to be 0, π, and 4π/3, or for a two-bin detection scheme, the phases can be selected to be 0 and 4π/3. - In the case of a two-bin detection scheme, rather than obtaining a quadrature amplitude from each
pixel 68, the power of thesignal light 44 can be computed as the absolute difference between the two intensity values stored in the two bins 70 for eachpixel 68 and then averaged in accordance with the following equation: - [7] Psignal∝Avg(Σk|Value1,k−Value2,k|) across all k=1, 2, . . . N speckles or pixels, where Value1 is the intensity value in the first bin 70 of the
respective pixel 68, and Value2 is the intensity value in the second bin 70 of therespective pixel 68. Thus, it can be appreciated that the intention of the two-bin measurement is to arrive at a precise estimate of power of the signal light Psignal up to the aforementioned scaling relationship defined by the strength of the reference light Preference, by removing the terms that are constant between the two measurements of the lock-incamera 28, and removing the unknown speckle-specific phases, and instead extracting only the amplitude of the cosine term. In the context of the UOT, just as with the quadrature detection scheme described above, the two-bin detection scheme serves as a measurement of light absorption at a single spatial voxel within the tissue. - However, the two-bin detection scheme represents a simplification that leads to only a small constant decrease factor in the signal to noise ratio. The dominant signal source contributing to detection has the following number of photons impinging on one pixel:
- [4] dominant signal=(ε/hv)×B×τ/2×|cos(0+φ)−cos(π+φ)|average over φ in [0,2π];
and the dominant noise source in the two-bin measurement of this amplitude is due to the shot noise in thereference light 42, and has the following number of photons impinging on one pixel: - [5] dominant noise=((ετ/hv)1/2×(2/π)B/A1/2,
where A and B are constants in the equation A+B(cosφ), B is proportional to the number of tagged photons perdetector 68, ε is a detector efficiency scaling factor, hv is the per-photon energy (with h as Plank's constant, and v as the frequency of the light), and τ is the integrated pulse widths used in the measurement, and φ is a random, pixel-specific speckle phase. - With a number of pixels N, the signal-to-noise ratio (SNR) scales with N1/2, since the total shot noise grows as N1/2, whereas the total signal grows as N, so that:
- [6] SNRN pixels=(N×(ε τ/hv)1/2×(2/π)×B/A1/2, which shows that, just as in the quadrature detection scheme, the SNR improves with increasing number of pixels in the lock-in camera 2828, and the Poisson nature of photon shot noise statistics is being utilized to determine the fundamental signal to noise ratio.
- Notably, the use of a two-bin detection scheme, rather than the four-bin quadrature scheme, provides the advantage that only two optical pulses, as opposed to four optical pulses, needs to be generated, thereby shortening the time period needed to take a measurement of the
target voxel 14, and thus, alleviating the speckle decorrelation time limitation. - In an optional embodiment, a digital optical phase conjugation (DOPC) technique can be used to boost the sensitivity of the pulsed UOT detection. DOPC can be performed in the context of schemes that rely on time reversal based optical phase conjugation using “guidestars” localized in three dimensions, for instance, using schemes, such as Time Reversal of Ultrasound-Encoded Light (TRUE) (see, e.g., Xu X, Liu H., Wang L V, “Time-Reversed Ultrasonically Encoded Optical Focusing into Scattering Media,” Nature Photonics, Vol. 5, No. 3, pp. 154-157 (Mar. 1, 2011); Wang Y M, Judkewitz B, DiMarzio C A, Yang C., “Deep-Tissue Focal Fluorescence Imaging with Digitally Time-Reversed Ultrasound-Encoded Light,” Nature Communications, Vol. 3, Article 928 (Jun. 16, 2012); Horstmeyer R., Ruan H, Yang C, “Guidestar-Assisted Wavefront-Shaping Methods for Focusing Light into Biological Tissue,” Nature Photonics, Vol. 9, No. 9, pp. 563-571 (Sep. 1, 2015).
- These methods are used to focus light to a guide-star-defined point deep inside a scattering medium, by measuring the wavefront emanating from the guidestar and digitally time-reversing (e.g., phase conjugating) light in order to cause the light to “play back” its path through the scattering medium and come to focus at the guidestar position. In the context of UOT, the guidestar is the focal point of an ultrasound beam. In these methods, the phase of a tagged light field originating from a given three-dimensional guidestar voxel in the brain is measured using demodulation and quadrature detection, and then an approximate phase-conjugate, i.e., approximate time-reverse light field, possibly amplified in intensity, is “played back” to focus light to the three-dimensional guidestar location despite the effects of strong or even diffusive scattering in the tissue.
- In the context of the
UOT system 10 described herein, the phase of the wavefront of thesignal light 44 originating from the target voxel 14 (i.e., the guidestar) is measured using the pulsed UOT detection scheme described above, as illustrated inFIG. 9 a, with the exception that, in addition to extracting the power of the wavefront of the Psignal, the unknown phase φunknown of the wavefront of thesignal light 44 is extracted using the known principles of quadrature detection. As illustrated inFIG. 9 b, the wavefront of thesignal light 44 is then amplified and retransmitted back in the opposite direction and focused onto the target voxel 14 (i.e., the guidestar), where it is tagged by theultrasound 32 and may then be detected by the same or another lock-in camera so as to perform the process iteratively. The retransmission should be timed such that the light through the voxel is coincident in time with an ultrasound pulse passing through the voxel. - Referring to
FIG. 10 a, the phase of thesignal light 44 extracted in the pulsed UOT detection scheme may be used in conjunction with a spatial light modulator (SLM)array 78 that is co-registered (e.g., pixel-by-pixel) with the lock-incamera 28 to perform optical phase conjugation of the signal light 44 (see, e.g., Laforest T, Verdant A, Dupret A, Gigan S., Ramaz F, Tessier G, “Co-Integration of a Smart CMOS Image Sensor and a Spatial Light Modulator for Real-Time Optical Phase Modulation,” Proc. Of SPIE-IS&T, Vol. 2014, 9022:90220N-1 (March 2014). - The
SLM array 78 may include any of a number of different amplitude and/or phase modulator structures, such as liquid crystals, re-positionable microelectromechanical systems (MEMS) mirrors, ferroelectrics, digital micro-mirror device pixels, among others. In one embodiment, theSLM array 78 may be semi-transparent (e.g., a liquid crystal modulator backed by a partial reflector), and can be inserted into the light path between the entry of thereference light 42 and the lock-incamera 28. TheSLM array 78 may be built directly on top of the lock-incamera 28 to create a phase conjugation array, with this arrangement being similar to the pulsed UOT detection scheme described above. - Referring to
FIG. 10 b, post-detection, eachpixel 68 of the lock-incamera 28 will send the conjugate phase information to the SLM array 78 (conjugate phase information being the negative of the detected phase). Each pixel may have internal electronics (e.g., transistors) that compute and sends the desired phase adjustment as an electrical voltage to adjust the corresponding pixel phase or amplitude of theSLM array 78. In a phase-only optical phase conjugation scheme, each pixel of the phase conjugation array will simply “actuate” the conjugate phase, such that light reflected from the pixel will accrue the phase. Amplitude-only phase conjugation can alternatively be performed by reflecting or not reflecting the input light based on the conjugate phase information. In aggregate, thesame reference light 42 used in the detection process (while blocking the sample light 40) or light precisely aligned with thereference light 42 will reflect off the phase conjugation array to create a phase conjugatelight field 80 that will focus back to thetarget voxel 14. - The improvement in contrast of this return light 80 (i.e., the phase conjugate light field) to the
target voxel 14 is given by: Contrast A=α*((N−1)/M+1), wherein N is the number of input optical modes (or the number of photons if less than the number of input optical modes), which is approximately equal to the number of pixels on the phase conjugation array); M is the number of target optical modes in thetarget voxel 14, and aequals 1 when a full phase and amplitude conjugation is performed, and is some value smaller than 1 when a phase only, amplitude only, and/or coarse grain phase conjugation is performed. The term “coarse grain,” in this context, means that the phase playback at each pixel can take only a finite number of possible values. - The phase conjugation process can be iterated many times, each time taking the light field, resulting from the last step, and phase conjugating that scattered light field. The contrast improvement can be expected to grow as (contrast A)K, where K is the number of iterations. Thus, the number of photons traveling through the
target voxel 14 can be exponentially amplified, thereby improving the effective modulation depth of the UOT (i.e., the fraction of the ultrasound tagged photons reaching the detector). The addition of phase conjugation to the pulsedUOT system 14 could be used to increase the number of collected tagged photons, increase modulation depth, or decrease ultrasound intensity or duty cycle requirements. - Performance estimates for the
UOT system 10 described herein in the detection of a blood-oxygen-level dependent signal in the brain through the skull as a function of the number of pixels in the lock-incamera 28 used (in this case, 10 million pixels or higher) indicate that neural activity dependent changes in the blood-oxygen-level dependent signal could be detected at hundreds to thousands of voxels per 100 millisecond temporal sample. In this calculation, the use of a 2 MHz ultrasound, and thus a spatial resolution on the order of ½ millimeter, is assumed, exceeding the spatial resolution of traditional blood-oxygen-level dependent signal measurements, like functional MRI (fMRI), and vastly exceeding the many millimeter to multiple centimeter-scale spatial resolution of diffuse optical tomography, including time-gated forms of diffuse optical tomography. In this calculation, it is further assumed that millions of tagged photons must be collected from thetarget voxel 14 per temporal sample in order to measure naturally occurring blood-oxygen-level dependent signals functional changes in the human brain, which are on the order of small fractions of a percent, while overcoming shot noise fluctuations in the number of detected tagged photons. - In one embodiment, the
processor 30 utilizes blood-oxygen-level dependent signals detected by the lock-incamera 28 to determine the neural activity in the brain; that is, blood-oxygen-level dependent signals provide a sense of the level of deoxygenated and/or oxygenated hemoglobin concentration or relative abundance in thetarget voxel 14 in the brain, and given the known coupling between cerebral hemodynamics and neuronal activity, theprocessor 30 can thereby determine the extent of neuronal activity in thattarget voxel 14. In another embodiment, theUOT system 10 detects blood-oxygen-level dependent signals over multiple wavelengths of the sample light, in which case, theprocessor 30 may determine and compare the optical absorption characteristics of thetarget voxel 14 of blood-oxygen-level dependent signals over the different wavelengths of sample light in order to determine the level of deoxygenated and/or oxygenated hemoglobin concentration or relative abundance present in thetarget voxel 14 according to known principles of functional infrared spectroscopy, for instance by solving two equations in two unknowns relating the measured absorption at two wavelengths to the level of deoxygenated and/or oxygenated hemoglobin concentration or relative abundance in the blood, or alternatively several equations in several unknowns representing absorption at several wavelengths in order to determine the concentrations of several molecular species in thetarget voxel 14. - In one particularly advantageous embodiment, instead of detecting blood-oxygen-level dependent signals, the
processor 30 may detect faster signals of neuronal activity, such as in the brain, to determine the extent of neuronal activity in thetarget voxel 14. Neuronal activity generates fast changes in optical properties, called “fast signals,” which have a latency of about 10-100 milliseconds and are much faster than the metabolic (approximately 100-1000 milliseconds) and hemodynamic (hundreds of milliseconds to seconds) evoked responses (see Franceschini, M A and Boas, D A, “Noninvasive Measurement of Neuronal Activity with Near-Infrared Optical Imaging,” Neuroimage, Vol. 21, No. 1, pp. 372-386 (January 2004)). Additionally, is believed that brain matter (e.g., neurons and the extracellular matrix around neurons) hydrates and dehydrates as neurons fire (due to ion transport in and out of the neurons), which could be measured via determining the absorption characteristics of water in thetarget voxel 14. In this case, it is preferred that thetarget voxel 14 be minimized as much as possible by selecting the appropriate ultrasound frequency (e.g., two to six times the size of a neuron, approximately 100 micrometers) in order to maximize sensitivity to highly localized changes in fast indicators of neural activity. As illustrated inFIG. 11 , the optical absorption coefficient of water is relatively high for wavelengths of light in the range of 950 nm-1080 nm. Thus, for maximum sensitivity to changes in optical absorption of tissue due to changes in the level of water concentration or relative water concentration in the brain matter, it is preferred the wavelength of the sample light be in the range of 950 nm-1080 nm. - Regardless of the nature of the detected signal and physiologically-dependent optical parameter, the
processor 30 may optionally use a computational model of light propagation in the tissue, and deconvolution or inverse problem optimization methods/algorithms, to improve the spatial resolution of the resulting measurement. Empirical measurements of a sample may be compared to those predicted by a model of the spatial layout of absorbers of the sample incorporating an effective point spread function of detection, such that the model may be improved to obtain an optimal match between the model predictions and the observed signals from the sample (see Powell S., Srridge S R, Leung T S, “Gradient-Based Quantitative Image Reconstruction in Ultrasound-Modulated Optical Tomography: First Harmonic Measurement Type in a Linearized Diffusion Formulation,” IEEE Transactions on Medical Imaging, Vol. 35, No. 2, pp. 456-467 (February 2016). - Although the
UOT system 10 has been described herein as acquiring only one measurement of thetarget voxel 14, it should be appreciated that theUOT system 10 may acquire multiple measurements of thetarget voxel 14 over time that yields a time trace indicative of time varying physiologically depending optical properties in thetarget voxel 14, such as time-varying optical absorption in thetarget voxel 14 due to functional changes in the brain. Optionally, two time traces of thetarget voxel 14 can be acquired, one time trace being generated with theultrasound 32 turned on at regular intervals in the same manner described above, and another time trace generated with theultrasound 32 turned off at regular intervals. For example, a measurement of thetarget voxel 14 may be acquired when theultrasound 32 turned on to create a first data point on the first time trace; a measurement of thetarget voxel 14 may be acquired when theultrasound 32 turned off to create a first data point on the second time trace; a measurement of thetarget voxel 14 may be acquired when theultrasound 32 turned on to create a second data point on the first time trace; a measurement of thetarget voxel 14 may be acquired when theultrasound 32 turned off to create a second data point on the second time trace; and so forth. The second time trace may provide a baseline null signal measurement trace, which is useful for tracking secondary variations distinct from the first time trace's signal variations due to theultrasound 32. - Referring now to
FIG. 12 , the physical implementation of theUOT system 10 will be described. As there shown, theUOT system 10 includes awearable unit 90 that is configured for being applied to thepatient 18, and in this case, worn on the head of thepatient 18, and an auxiliary head-worn or not head-wornunit 92 coupled to thewearable unit 90 via a wired connection 94 (e.g., electrical wires). Alternatively, theUOT system 10 may use a non-wired connection (e.g., wireless radio frequency (RF) signals) for providing power to or communicating between components of the respectivewearable unit 90 andauxiliary unit 92. - In the illustrated embodiment, the
wearable unit 90 includes asupport structure 94 that either contains or carries thetransducer arrangement 34 of the acoustic assembly 20 (shown inFIG. 2 ), theinterferometer 22, and the lock-incamera 28. Thewearable unit 90 may also include anoutput port 98 a from which theultrasound 32 generated by the acoustic assembly 20 (shown inFIG. 1 ) is emitted, anoutput port 98 b from which thesample light 38 generated by the interferometer 22 (shown inFIG. 1 ) is emitted, and aninput port 98 c into which thesample light pattern 47 comprising the tagged signal light and untagged background light are input into theinterferometer 22. It should be appreciated that although theinput port 98 c is illustrated in close proximity to theinput ports input port 98 c and theoutput ports support structure 94 may be shaped, e.g., have a banana, headband or hat shape, such that the ports 98 are in close contact with the outer skin of the body part, and in this case, the scalp of thehead 18 of thepatient 18. An index matching fluid maybe used to reduce reflection of the light generated by thelight source 30 of theinterferometer 22 from the outer skin, and an ultrasound transmitting adhesive or acoustic coupling material can be used to facilitate conduction of theultrasound 32 into the body part. An adhesive or belt (not shown) can be used to secure thesupport structure 94 to the body part. - The
auxiliary unit 92 includes ahousing 96 that contains thecontroller 24 and the processor 30 (shown inFIG. 1 ). Theauxiliary unit 92 may additionally include a power supply (which if head-worn, may take the form of a rechargeable or non-chargeable battery), a control panel with input/output functions, a display, and memory. - The
auxiliary unit 92 may further include thesignal generator 36 of theacoustic assembly 20, as well as any drive circuitry used to operate theinterferometer 22. - The
interferometer 22 and lock-incamera 28 are preferably mechanically and electrically isolated from theacoustic assembly 20, such that the emission of theultrasound 32 by theacoustic assembly 20, as well as the generation of RF and other electronic signals by theacoustic assembly 20 minimally affects the detection of the optical signals by theinterferometer 22 and generation of data by the lock-incamera 28. Thewearable unit 90 may include shielding (not shown) to prevent electrical interference and appropriate materials that attenuate the propagation of acoustic waves through thesupport structure 94. - Having described the arrangement of function of the
UOT system 10, one method of operating the UOT system on a patient will now be described. In this method,ultrasound 32 is delivered into thetarget voxel 14 in theanatomical structure 16, and sample light 40 is delivered into theanatomical structure 16, wherein aportion 40 a of thesample light 40 passing through thetarget voxel 14 is scattered by theanatomical structure 16 as thesignal light 44, and anotherportion 40 b of thesample light 40 not passing through thetarget voxel 14 is scattered by theanatomical structure 16 asbackground light 46 that combines with thesignal light 44 to create thesample light pattern 47. As exemplified above, theanatomical structure 16 may be an intact head comprising the scalp, skull, and brain matter. Due to the high resolution of theUOT system 10, thetarget voxel 14 may be smaller than one mm3. - The
reference light 42 is combined with thesample light pattern 47 to generate an interference light pattern 48 (e.g., in a homodyne manner), and in this method, a speckle light pattern. Theultrasound 32 and sample light 40 are pulsed in synchrony, such that only thesignal light 44 is shifted (i.e., tagged) by theultrasound 32. That is, as described above, each pulse of thesample light 40 will pass through thetarget voxel 14 only as theultrasound 32 passes through thetarget voxel 14, such that no portion of thebackground light 46 will be tagged by theultrasound 32. Theinterference light pattern 48 is sequentially modulated to generate a plurality of different interferencelight patterns 48. The spatial components of any particular interferencelight pattern 48 can then be simultaneously detected, and a plurality of values can be stored in the respective bins 70 (either inbins 70 a, inbins 70 b, inbins 70 c, orbins 70 d) of thedetectors 68. The values are representative of the spatial component for the respective interferencelight pattern 48. The physiologically-dependent optical parameter of thetarget voxel 14 is then determined based on the spatial component values stored in the bins 70. Due to the high speed of the lock-incamera 28, the spatial components for any particular interferencelight pattern 48 may be simultaneously detected and stored in the respective bins 70 very quickly. For example, in one embodiment, the spatial components for any particular interferencelight pattern 48 may be simultaneously detected, and the resulting spatial component values for all theinterference light patterns 48 are stored in the respective bins 70 within 1 millisecond. In another embodiment, the spatial components for any particular interferencelight pattern 48 may be simultaneously detected, and the resulting spatial component values for all theinterference light patterns 48 are stored in the respective bins 70 within 1 microsecond to 1 millisecond. - Referring to
FIG. 13 , oneparticular method 100 performed by theUOT system 10 to non-invasively image thetarget voxel 14 inanatomical structure 16 will now be described. Thisparticular method 100 implements the pulsing sequence ofFIG. 5 . - The
controller 24 operates theacoustic assembly 20 to generate and deliver a pulse ofultrasound 32 having a frequency fus (initially, ultrasound pulse Ua illustrated inFIG. 5 ) into theanatomical structure 16, e.g., by sending a control signal to thesignal generator 36 to pulse an electrical signal on and off (step 102). Thecontroller 24 sets the phase difference between thesample light 40 and thereference light 42 to one of the four pre-defined values (0, π/2, π, and 3π/2) by sending a control signal to thephase shifter 54 of the interferometer 22 (step 104). This pre-defined phase difference value may be first set to 0. Next, thecontroller 24 operates theinterferometer 22 to generate and emit a pulse of source light 38 having a frequency f, e.g., by sending a control signal to the drive circuit to pulse thelight source 50 on and off (step 106). The interferometer 22 (e.g., via the beam splitter 52) splits the pulse of source light 38 into a pulse of sample light 40 (initially, the sample light pulse La illustrated inFIG. 5 ) and a pulse of reference light 42 (step 108). - The wavelength (and thus, the frequency f) of the source light 38 may be selected based on the physiologically-dependent optical parameter to be ultimately determined. For example, if the physiologically-dependent optical parameter is the level of deoxygenated and/or oxygenated hemoglobin concentration or relative abundance, the wavelength of the source light 38 may be in the range of 605 nanometers to 950 nanometers, whereas if the physiologically-dependent optical parameter to be determined is a water absorption level (level of water concentration or relative water concentration), the wavelength of the source light 38 may be in the range of 950-1080 nanometers.
- Next, prior to the pulse of sample light 40 entering the
anatomical structure 16, thecontroller 24 operates theinterferometer 22 to frequency shift the pulse ofsample light 40 by the ultrasound frequency fus, e.g., by sending a control signal to thefrequency shifter 56, resulting in the pulse of sample light 40 having a frequency f−fus (step 110). The frequency-shifted pulse ofsample light 40 is then delivered into and diffusively scattered within the anatomical structure 16 (step 112). As the pulse of frequency shiftedsample light 40 scatters diffusively through theanatomical structure 16, a portion will pass through thetarget voxel 14 and be frequency shifted (i.e., tagged) back to its original frequency f by the pulse ofultrasound 32 passing through thetarget voxel 14, resulting in a pulse ofscattered signal light 44 having the same frequency f (step 114); and remaining portion will not pass through thetarget voxel 14, and thus will not be frequency shifted by the pulse ofultrasound 32, resulting in a pulse ofscattered background light 46 having a frequency f−fus (the same frequency as the frequency shiftedsample light 40 prior to entering the anatomical structure 16) (step 116). - Next, the
interferometer 22 then combines (e.g., via the light combiner 58) the pulse of reference light 42 with the pulses ofsample light pattern 47 to generate a pulse of an interference light pattern 48 (initially, the interference light pattern pulse Ia illustrated inFIG. 5 ) (step 118). Then, under control of thecontroller 24, all of the detectors 68 (FIG. 4 ) of the lock-incamera 28 simultaneously detect respective spatial components of the interference light pattern 48 (i.e., speckle grains in the case where the interference light pattern includes a speckle light pattern) (step 120), and values (e.g., power values) representative of the spatial components of theinterference light pattern 48 are stored in bins 70 (initially, thefirst bins 70 a of the corresponding detectors 68) (step 122). - At this point, only one quadrature measurement has been taken. If the
interferometer 22 has not been set to all four of the phases (step 124), thecontroller 24 then repeats steps 102-122 to take the next quadrature measurement. That is, the next pulse of ultrasound 32 (e.g., ultrasound pulse Ub illustrated inFIG. 5 ) is generated and emitted into the anatomical structure 16 (step 102); the phase difference between the sample light 40 and the reference light 42 is set to the next pre-defined value (e.g., 7/2) (step 104); the next pulse of source light 38 is generated (step 106) and split into the next pulse of sample light 40 (e.g., sample light pulse Lb illustrated inFIG. 5 ) and a pulse of reference light 42 (step 108); the next pulse of sample light 40 is frequency shifted by the ultrasound frequency fus (step 110); the frequency shifted pulse of sample light 40 is delivered and diffusively scattered within the anatomical structure 16 (step 112); a portion of the scattered pulse of sample light 40 passing through the target voxel 14 is frequency shifted (tagged) by the pulse of ultrasound 32 passing through the target voxel 14, thereby generating the next pulse of scattered signal light 44 (step 114); the remaining portion of the scattered pulse of sample light 40 not passing through the target voxel 14 is not frequency shifted (not tagged) by the pulse of ultrasound 32 passing through the target voxel 14, thereby generating the next pulse of scattered background light 46 (step 116); the next pulses of reference light 42 and sample light pattern 47 are combined into the pulse of the next interference light pattern 48 (e.g., the interference light pattern pulse Ib illustrated inFIG. 5 ) (step 118); the spatial components of the next pulse of the interference light pattern 48 are detected (step 120); and the resulting spatial component values are stored in the bins 70 (e.g., the second bins 70 b of the corresponding detectors 68) (step 122). - Thus, it can be appreciated that steps 102-122 will be repeated to take the remaining quadrature measurements to generate and detect the pulses of the remaining interference light patterns (e.g., the third and fourth interference light pattern pulses Ic, Id illustrated in
FIG. 5 ) for the remaining phase settings (e.g., π and 3π/2) and ultimate storage of the spatial component values in the bins 70 (e.g., the third andfourth bins - After all four quadrature measurements have been taken, the
controller 24 recalls the spatial component values of the detected interferencelight pattern pulses 48 from the bins 70 of the lock-incamera 28 and transfers these values to the processor 30 (step 126). Theprocessor 30 reconstructs the amplitude of the signal light 44 from the fourinterference light patterns 48 based on these spatial component values (e.g., by using the quadrature equation [2]) (step 128). Steps 102-128 can be iterated to repeatedly acquire data measurements of thetarget voxel 14, and if a sufficient number of data measurements have been acquired (step 130), theprocessor 30 may then determine the physiologically-dependent optical parameter (e.g., level of deoxygenated and/or oxygenated hemoglobin concentration or relative abundance or level of water concentration or relative water concentration) of thetarget voxel 14 based on the data measurements (step 132). In the case where thetarget voxel 14 is brain matter, theprocessor 30 may further determine the level of neural activity within thetarget voxel 14 based on the determined physiologically-dependent optical parameter (step 134). - For example, if the physiologically-dependent optical parameter is the level of deoxygenated and/or oxygenated hemoglobin concentration or relative abundance, and if the amplitude of the
signal light 44 is relatively low (or high), indicating high absorption of light by blood in thetarget voxel 14, it can be assumed that there is a relatively high (or low) hemodynamic response (depending on the light wavelength used) through thetarget voxel 14, and thus, a substantial amount of neural activity in thetarget voxel 14. In contrast, if the amplitude of thesignal light 44 is relatively high (or low), indicating low absorption of light by blood in thetarget voxel 14, it can be assumed that there is a relatively low hemodynamic response (depending on the wavelength) through thetarget voxel 14, and thus, comparatively little neural activity in thetarget voxel 14. - If the physiologically-dependent optical parameter is the level of water concentration or relative water concentration, and if the amplitude of the
signal light 44 greatly varies over a short period of time, indicating a fast signal of neural activity in the brain tissue, it can be assumed that there is a substantial amount of neural activity in thetarget voxel 14. In contrast, if the amplitude of thesignal light 44 varies very little over a short period of time, indicating that there is no fast signal of neural activity in the brain matter, it can be assumed that there is very little or no neural activity in thetarget voxel 14. - Referring to
FIG. 14 , anotherparticular method 100′ performed by theUOT system 10 to non-invasively image thetarget voxel 14 in theanatomical structure 16 will now be described. Thisparticular method 100′ implements the pulsing sequence ofFIG. 7 . - The
method 100′ is similar to themethod 100 illustrated inFIG. 13 , with the exception that, instead of a one-to-one correspondence between the sample light pulse and the ultrasound pulse, multiple sample pulses (in this case, four) are delivered to the biological tissue for every ultrasound pulse delivered to the biological tissue. Thus, after a quadrature measurement, the process returns to step 104 (instead of step 102) to take the next quadrature measurement. - That is, during the delivery of current pulse of ultrasound 32 (e.g., ultrasound pulse U illustrated in
FIG. 5 ), the phase difference between the sample light 40 and the reference light 42 is set to the next pre-defined value (e.g., 7/2) (step 104); the next pulse of source light 38 is generated (step 106) and split into the next pulse of sample light 40 (e.g., sample light pulse Lb illustrated inFIG. 5 ) and a pulse of reference light 42 (step 108); the next pulse of sample light 40 is frequency shifted by the ultrasound frequency fus (step 110); the frequency shifted pulse of sample light 40 is delivered and diffusively scattered within the anatomical structure 16 (step 112); a portion of the scattered pulse of sample light 40 passing through the target voxel 14 is frequency shifted (tagged) by the pulse of ultrasound 32 passing through the target voxel 14, thereby generating the next pulse of scattered signal light 44 (step 114); the remaining portion of the scattered pulse of sample light 40 not passing through the target voxel 14 is not frequency shifted (not tagged) by the pulse of ultrasound 32 passing through the target voxel 14, thereby generating the next pulse of scattered background light 46 (step 116); the next pulses of reference light 42 and sample light pattern 47 are combined into the pulse of the next interference light pattern 48 (e.g., the interference light pattern pulse Ib illustrated inFIG. 5 ) (step 118); the spatial components of the next pulse of the interference light pattern 48 are detected (step 120); and the resulting spatial component values are stored in the bins 70 (e.g., the second bins 70 of the corresponding detectors 68) (step 122). - After all four quadrature measurements have been taken at steps 102-124, as in the manner described above with respect to the
method 100 ofFIG. 13 , the spatial component values of the detected interferencelight pattern pulses 48 are recalled from the bins 70 of the lock-in camera 28 (step 126); and the amplitude of thesignal light 44 is reconstructed from the fourinterference light patterns 48 based on these spatial component values (step 128). Steps 102-128 can be iterated to repeatedly acquire data measurements of thetarget voxel 14, and if a sufficient number of data measurements have been acquired (step 130), the physiologically-dependent optical parameter (e.g., the level of deoxygenated and/or oxygenated hemoglobin concentration or relative abundance or level of water concentration or relative water concentration) of thetarget voxel 14 is determined based on the data measurements (step 132), and the level of neural activity within thetarget voxel 14 is determined based on the determined physiologically-dependent optical parameter (step 134). - Although particular embodiments of the present inventions have been shown and described, it will be understood that it is not intended to limit the present inventions to the preferred embodiments, and it will be obvious to those skilled in the art that various changes and modifications may be made without departing from the spirit and scope of the present inventions. Thus, the present inventions are intended to cover alternatives, modifications, and equivalents, which may be included within the spirit and scope of the present inventions as defined by the claims.
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