Fundus camera-guided ultrasound system and method for precise ocular lesion location and assessment

Abstract

Various examples are provided related to hybrid light-ultrasound fundus imaging. In one example, an ocular imaging system includes a substantially transparent ultrasound transducer; an optical imaging system including a fundus camera sensor aligned with the ultrasound transducer; and one or more control interface components that can enable visualization of an ocular surface and perform filtering on imaging signals generated by the ultrasound transducer operatively connected to the fundus camera sensor. In another example, a method for ocular imaging includes illuminating an ocular surface of an eye with the ocular imaging system; aligning the ultrasound transducer with a lesion on the ocular surface using the optical imaging system; and obtaining ultrasound pulse-echo scan data of the lesion using the ultrasound transducer.

Description

Docket: 320903-2510 FUNDUS CAMERA-GUIDED ULTRASOUND SYSTEM AND METHOD FOR PRECISE OCULAR LESION LOCATION AND ASSESSMENT CROSS REFERENCE TO RELATED APPLICATIONS

[0001] This application claims priority to, and the benefit of, U.S. provisional application entitled “Fundus Camera-Guided Ultrasound System and Method for Precise Ocular Lesion Location and Assessment” having serial no.63 / 551,472, filed February 8, 2024, which is hereby incorporated by reference in its entirety. STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

[0002] This invention was made with government support under grant R01EY028662, R01EY030126, R01EY035084, R01EY029673, R44EY028786, and P30EY001792 awarded by the National Institutes of Health. The Government has certain rights in the invention. This invention was also supported by the Richard and Loan Hill Department of Biomedical Engineering, University of Illinois Chicago and Research to Prevent Blindness. BACKGROUND

[0003] Reliable diagnosis of ocular lesions is an essential step in the management of many vision- and life-threatening ocular diseases. Light fundus photography provides valuable information for ocular disease diagnosis, lesion monitoring, and treatment response assessment. However, light-based examination has limited penetration capability for quantitative evaluation of thick lesions such as uveal melanoma or choroidal hemangiomas. Moreover, ocular media opacities such as corneal pathology, dense cataract, vitreous hemorrhage, or vitritis, may limit the light fundus imaging performance for reliable evaluation of ocular lesions. SUMMARY

[0004] Aspects of the present disclosure are related to hybrid light-ultrasound fundus imaging. In one aspect, among others, an ocular imaging system comprises: a substantially transparent ultrasound transducer; an optical imaging system comprising a fundus camera sensor aligned with the ultrasound transducer; and one or more control interface components configured to enable visualization of an ocular surface andDocket: 320903-2510 perform filtering on imaging signals generated by the ultrasound transducer operatively connected to the fundus camera sensor. In one or more aspects, the optical imaging system can comprise a miniature indirect illumination source operatively connectable to the one or more control interface components. The miniature indirect illumination source can comprise a source of light positioned to illuminate the ocular surface via one or more lens of the optical imaging system. The source of light can comprise a combination of broadband and narrowband light emitting diodes. The source of light can be configured to provide light at wavelengths greater than 500 nm. The source of light can be in a plane with and offset from a camera lens of the fundus camera sensor. The source of light can illuminate the ocular surface via the substantially transparent ultrasound transducer.

[0005] In various aspects, the substantially transparent ultrasound transducer can be a single ultrasound transducer. The substantially transparent ultrasound transducer can comprise a contact lens disposed across a side of the single ultrasound transducer adjacent to the fundus camera sensor. The single ultrasound transducer can be curved. The substantially transparent ultrasound transducer can comprise a contact lens disposed within a hollow middle portion of the single ultrasound transducer. In some or all aspects, the ocular imaging system can comprise a second substantially transparent ultrasound transducer and a contact lens disposed across both ultrasound transducers. The substantially transparent ultrasound transducer can comprise a lithium niobate (LNO) crystal with indium tin oxide (ITO) electrodes disposed on opposite sides of the LNO crystal. The substantially transparent ultrasound transducer can comprise a matching layer disposed on a side opposite the fundus camera sensor. The ocular imaging system can be a portable handheld imaging system.

[0006] In another aspect, a method for ocular imaging comprises: illuminating an ocular surface of an eye with the ocular imaging system; aligning the ultrasound transducer with a lesion on the ocular surface using the optical imaging system; and obtaining ultrasound pulse-echo scan data of the lesion using the ultrasound transducer. In one or more aspects, the method can comprise obtaining ultrasound pulse-echo scan data at a periphery of an ocular image using a second substantially transparent ultrasound transducer. The ocular surface can be illuminated with light at wavelengths greater than 500 nm. An illumination efficiency can be configurable to accurately guide the ultrasound transducer to the lesion on the ocular surface located in a posterior region of the eye.

[0007] Other systems, methods, features, and advantages of the present disclosure will be or become apparent to one with skill in the art upon examination of the following drawings and detailed description. It is intended that all such additionalDocket: 320903-2510 systems, methods, features, and advantages be included within this description, be within the scope of the present disclosure, and be protected by the accompanying claims. In addition, all optional and preferred features and modifications of the described embodiments are usable in all aspects of the disclosure taught herein. Furthermore, the individual features of the dependent claims, as well as all optional and preferred features and modifications of the described embodiments are combinable and interchangeable with one another. BRIEF DESCRIPTION OF THE DRAWINGS

[0008] Many aspects of the present disclosure can be better understood with reference to the following drawings. The components in the drawings are not necessarily to scale, emphasis instead being placed upon clearly illustrating the principles of the present disclosure. Moreover, in the drawings, like reference numerals designate corresponding parts throughout the several views.

[0009] FIG.1A is a schematic diagram illustrating an example of a transparent transducer, in accordance with various embodiments of the present disclosure.

[0010] FIG.1B illustrates the transparent transducer of FIG.1A with connections including a ground (G) and wire (W), in accordance with various embodiments of the present disclosure.

[0011] FIG.2A schematically illustrates an experimental setup with RCP indicating the retina conjugate plane, in accordance with various embodiments of the present disclosure.

[0012] FIG.2B is an image of an example of an imaging system, in accordance with various embodiments of the present disclosure.

[0013] FIGS.3A and 3B illustrate examples of simulated and measured electrical impedance and phase spectrum of the transducer, in accordance with various embodiments of the present disclosure.

[0014] FIGS.3C and 3D illustrate examples of simulated and measured pulse-echo performance of the transducer, in accordance with various embodiments of the present disclosure.

[0015] FIG.4A is an example of a fundus image taken from in vivo imaging on a rat eye, in accordance with various embodiments of the present disclosure.

[0016] FIG.4B illustrates an example of ultrasound pulse-echo data gathered from the same rat eye, in accordance with various embodiments of the present disclosure.Docket: 320903-2510

[0017] FIG.5A illustrates an example of the transmission spectrum of the transparent transducer, in accordance with various embodiments of the present disclosure.

[0018] FIGS.5B and 5C are examples of images of a retina of an eye model taken with the prototype and taken after replacing the ultrasound transducer in the prototype with a glass window, in accordance with various embodiments of the present disclosure.

[0019] FIG.6 conceptually illustrates an example of the proposed hybrid light- ultrasound imager, in accordance with various embodiments of the present disclosure.

[0020] FIG.7A illustrates a front view of an example of a two-element ultrasound probe for fundus photography guided A-scan measurement, in accordance with various embodiments of the present disclosure.

[0021] FIG.7B illustrates a cross-section view of an example of a linear array ultrasound probe for fundus photography guided B-scan measurement, in accordance with various embodiments of the present disclosure.

[0022] FIG.8 illustrates a design cross-section of an example of the proposed transparent ultrasonic transducer, in accordance with various embodiments of the present disclosure.

[0023] FIG.9 conceptually illustrates an example of the proposed hybrid light- ultrasound imager, in accordance with various embodiments of the present disclosure.

[0024] FIG.10A is a cross-section schematically illustrating an example of a transparent ultrasonic transducer, in accordance with various embodiments of the present disclosure.

[0025] FIG.10B is an image of an example of the transparent ultrasonic transducer of FIG.10A, in accordance with various embodiments of the present disclosure.

[0026] FIG.10C illustrates an example of measurement of optical transmission through the fabricated transparent ultrasonic transducer of FIG.10B, in accordance with various embodiments of the present disclosure.

[0027] FIG.11 is a cross-section schematically illustrating an example of a contact lens based transparent ultrasonic transducer, in accordance with various embodiments of the present disclosure.

[0028] FIGS.12A and 12B illustrate examples of simulated pulse-echo performance with bandwidth, electrical impedance magnitude and phase spectrum of the contact lens based transparent ultrasonic transducer, in accordance with various embodiments of the present disclosure.

[0029] FIG.13A is a cross-section schematically illustrating an example of a transparent ultrasonic transducer in ring-type form, in accordance with various embodiments of the present disclosure.Docket: 320903-2510

[0030] FIGS.13B and 13C illustrate examples of simulated pulse-echo performance with bandwidth, electrical impedance magnitude and phase spectrum of the transducer in ring-type form, in accordance with various embodiments of the present disclosure.

[0031] FIGS.14A and 14B illustrate an example of a proposed portable fundus camera with miniaturized indirect illumination, in accordance with various embodiments of the present disclosure.

[0032] FIG.14C includes examples of fundus images obtained with the fundus camera of FIGS.14A-14B, in accordance with various embodiments of the present disclosure.

[0033] FIGS.15A and 15B illustrate an example of the optical imaging system, in accordance with various embodiments of the present disclosure.

[0034] FIG.16 illustrates examples of images of a full fundus of a young adult, an a model eye with metal markers, and NIR imaging of the fundus, in accordance with various embodiments of the present disclosure.

[0035] FIG.17 illustrates examples of conventional widefield light fundus imaging and ultrasound and the proposed hybrid light-ultrasound imaging and ultrasound, in accordance with various embodiments of the present disclosure. DETAILED DESCRIPTION

[0036] Disclosed herein are various examples related to hybrid light-ultrasound fundus imaging. Reference will now be made in detail to the description of the embodiments as illustrated in the drawings, wherein like reference numbers indicate like parts throughout the several views.

[0037] A hybrid light-ultrasound fundus imager is presented that can enable rapid detection and accurate quantitative assessment of ocular lesions. Ultrasound imaging complements some limitations of light fundus imaging. However, its usability may be limited by difficulty to accurately localize the ultrasound probe to a lesion of interest. Therefore, there is an unmet need for developing a hybrid light-ultrasound fundus imager to foster accurate localization for quantitative assessment of focal ocular lesions.

[0038] Choroidal nevi are benign ocular tumors that are estimated to be present in 5% of adults over 40 years of age in the United States. A small fraction of these lesions transform into malignant lesions called uveal melanoma, which affect 6 non-Hispanic white individuals per million and at a lower rate in other races. However, due to being situated in the delicate structures of the eye, intraocular tumor biopsy and subsequent pathological examination are performed for diagnostic purposes only in exceptional casesDocket: 320903-2510 where the ocular oncologist is unable to make a conclusive diagnosis after clinical examination and systemic evaluation.

[0039] After diagnosis of uveal melanoma, patients typically undergo enucleation or globe sparing radiation therapy that can severely affect the vision. Early diagnosis of these malignant lesions can avoid enucleation and significantly preserve the patient's eyesight. Furthermore, even with successful local disease control, 50% of the patients diagnosed with uveal melanoma develop metastatic disease, with a high rate of subsequent mortality. Therefore, early detection of malignancy is of paramount importance to preserve eyesight and improve survival in patients with uveal melanoma.

[0040] The evaluation of a suspected ocular tumor can be performed with dilated fundus examination and multi-modal imaging to accurately identify benign and malignant lesions including choroidal nevi, uveal melanoma, choroidal metastasis, circumscribed choroidal hemangioma, choroidal osteoma, and retinoblastoma.

[0041] With the advent of widefield and ultra-widefield fundus cameras, identifying suspected lesions even in the periphery of the retina is possible. Fundus images provide important information, such as the size, location, color, presence of orange pigment, and presence of overlying drusen. However, traditional fundus photography lacks depth- resolved information, and typically shows the superficial layer information which is not enough to determine other important characteristics of an ocular tumor. Spectral-domain optical coherence tomography with enhanced depth-resolved imaging (EDI SD-OCT) capability has shown promise to accurately illustrate the intrinsic optical characteristics of ocular lesions.

[0042] Despite being able to distinguish lesions from normal tissue, the low penetration capability of OCT limits its usage to only very thin lesions, typically less than 1 mm in thickness, especially in melanocytic lesions. Imaging techniques such as fluorescein angiography (FA), indocyanine green angiography (ICGA), and fundus autofluorescence (FAF) can provide valuable information, such as changes in the retinal pigmented epithelium (RPE) overlying ocular lesions, tumor vasculature, and associated vascular leakage. However, these imaging modalities do not provide sufficient information related to the tumor depth, and their standalone diagnostic accuracy is relatively low.

[0043] Ultrasound imaging of the eye has been a longstanding imaging technique due to its excellent tissue penetration capability, non-invasive nature, and ability to differentiate tissues with different echogenicity. Experimental research with different ultrasound transducer materials and frequencies is also gaining popularity for non- invasive imaging of the eye. Usage of ultrasound is important in many subspecialities of ophthalmology including cataract surgery planning, glaucoma, vitreoretinal surgery,Docket: 320903-2510 uveitis, oculoplastic procedures, and emergency medicine in addition to ocular oncology where it serves the purpose of diagnosing and monitoring ocular tumors, such as uveal melanoma.

[0044] For imaging the posterior retina, both a one-dimensional amplitude scan (A- scan) and a two-dimensional brightness scan (B-scan) can be used. The B-scan can be used to observe anterior and posterior segments of the ocular lesion as well as the axial and lateral dimensions of the eye. Although a one-dimensional vector A-scan can be derived from a B-scan image, its usability for quantitative measurement of lesion thickness is limited, as the logarithmic amplification used to generate the B-scan image prevents the quantitative analysis needed for accurate lesion echogenicity analysis. The logarithmic amplification increases the dynamic range, however, reduces the sensitivity, which further prevents the differentiation between tissues where the variation of echogenicity is minimal.

[0045] Moreover, the focused beam used in the B-scan probe further complicates the process of computational echogenicity analysis of ocular lesions. A specifically designed A-scan probe called standardized A-scan with non-focused parallel beam and S-shaped amplification was optimized for eye bio-microscopy and has been an important part of many ophthalmology clinics ever since. Standardized A-scan has become the gold standard for calculating the apical height of ocular lesions, specifically because of both its enhanced penetration capability and its ability to differentiate the lesion border from the scleral tissue. Several quantifiable parameters other than the apical height have been extracted out of A-scan ultrasound, such as median internal reflectivity (%) of a lesion, heterogenous / homogenous echogenicity, number of internal reflectivity peaks inside a lesion, sound attenuation pattern (angle ^^) to facilitate an accurate clinical diagnosis.

[0046] The internal reflectivity and sound attenuation patterns are well established for differentiating choroidal melanoma from other lesions such as choroidal nevus, retinoblastoma, circumscribed choroidal hemangioma, choroidal metastasis. Given that treated choroidal melanomas show a considerable rise in median internal reflectivity, the A-scan ultrasound can also be useful in assessing the effectiveness of treatment.

[0047] Despite the promise, ultrasound A-scan comes with a set of challenges. Precise guidance of the A-scan line towards the lesion needs an experienced operator, especially when the lesion is relatively small. Inaccurate guidance could lead to erroneous measurement of parameters such as apical height and internal reflectivity. Furthermore, the growth of benign lesions, such as choroidal nevi, for example, can indicate its transformation towards malignant choroidal melanoma, and clinicians often observe the nevus size over a long period of time. However, monitoring the apical heightDocket: 320903-2510 from the exact same spot over a prolonged period is nearly impossible because of the absence of A-scan guidance. Therefore, accurate guidance an of A-scan ultrasound probe is highly desirable for quantitative analysis, management, and documentation of ocular lesions.

[0048] In accordance with the principles of the present disclosure, the feasibility of using a fundus camera to accurately guide an ultrasound A-scan probe in the posterior region of the eye is demonstrated. Further control and signal enhancement components are integrated into the fundus-guided ultra- sound system.

[0049] In an exemplary embodiment, a widefield fundus camera integrated with a transparent LiNbO3single-crystal ultrasound transducer facilitated the guidance of ultrasound localization. In vivo experiments with rat eyes demonstrated the ability of the prototype to simultaneously gather fundus images and A-scans of the eye. The resultant A-scan clearly differentiated the anterior cornea, the eye lens, and the posterior retina. The optical performance of the imager was also examined with a model human eye and compared with the performance of the same imager but with a glass window replacing the transparent transducer. The resultant images show that the imager with a transparent transducer can clearly visualize the retina and can be readily implemented in clinical settings. Materials and Methods

[0050] Design of the transparent ultrasound probe. Based on the transparent application for the transducer, a transparent lithium niobate (LNO) single crystal was selected as the core component for the ultrasound transducer fabrication. The acoustic impedance of the LNO is 34 MRayl. A Krimboltz, Leedom, and Mattaei (KLM) transducer equivalent circuit model-based modeling software PiezoCAD (Sonic Concepts, Inc., Woodinville, WA, USA) was applied to simulate and optimize the electrical impedance and pulse-echo performance of the designed transducer (see FIGS.3A-3D). To ensure the transparency for light penetration, a colorless parylene film (e.g., Parylene C, Specialty Coating Systems Inc., Indianapolis, IN, USA, 2023) with an acoustic impedance of 2.5 MRayl was chosen as the matching layer for an acoustic impedance compensation, and a transparent epoxy (e.g., EPO-TEK 301, Epoxy Technology, Inc., Billerica, MA, USA, 2023) with an acoustic impedance of 3.05 MRayl was determined as the backing layer for absorbing reflected ultrasound and penetrating optical source. The designed parameters of the transducer are summarized and listed in Table 1. Table 1. Design parameters of the transparent ultrasound transducer. Parameter ValueDocket: 320903-2510 Center frequency 40 MHz Surface area 5mm x 5mm [005nsparent LNO disk wafer (Boston Piezo-Optics, Inc., Bellingham, MA, USA) with a size of 5 mm x 5 mm was purchased and manufactured for the exemplary system. Both sides of the LNO were sputtered indium tin oxide (ITO), a transparent electrode, by the sputtering system (e.g., NSC-3000 Sputter Coater, Nano-Master, Inc., Austin, TX, USA). Afterward, the LNO with electrodes was shielded by the brass housing and connected with one side of the double-shield coaxial cable, and the other side of the cable was connected with the Sub-Miniature version A (SMA) connector for ultrasound driving system connection.

[0052] Degassed epoxy (e.g., EPO-TEK 301) was poured into the brass housing to form the backing layer. After 3 hours of curing for the epoxy, the Au / Cr electrode (100 nm I 50 nm) was sputtered to conduct the ITO edge of the LNO and brass housing for ground connection. Finally, a parylene film was coated on the surface of the LNO as a matching layer and water-proof layer. The fabrication of the transparent transducer was then finished and further characterized and tested. FIG.1A is a schematic diagram illustrating an example of the transparent ultrasound probe 103. The transducer has 5 mm X 5 mm of clear penetration window. FIG.1B shows a finished cylinder-type transparent ultrasound transducer. The letters printed on the background can easily be read, further indicating the transparency of the transducer.

[0053] Testing the performance of the transparent ultrasound probe. The electrical impedance and phase spectrum of the fabricated transparent transducer were measured by an impedance analyzer (e.g., HP 4294A, Agilent Technologies Inc., Santa Clara, CA, USA). The pulse-echo performance of the transparent transducer was measured by a set-up detailed in previous work. The transducer was immersed into a water tank, and the quartz plate was applied as a reflector for reflecting generated ultrasound. The transducer excited by pulser-receiver (e.g., Panametrics 5900PR, Olympus NDT Inc., Waltham, MA, USA) with 1 µJ energy per pulse trans- mitted and received the ultrasound. The axial resolution (^^^௫^^^) of the transducer can be calculated using the following equation (1).Docket: 320903-2510 ^^^௫^^^ ൌ ^ఒ (1) ଶ^^ where ^^ is the wavelength of the -6 dB fractionalspectrum of the transparent transducer was calculated by passing the light from a broadband LED (e.g., MBB1L3, Thorlabs Inc., Newton, NJ, USA) through the transducer and analyzing the transmitted light using a UV-VIS spectrometer (e.g., USB400, Ocean Optics Inc., Dunedin, FL, USA).

[0054] Experimental setup. FIG.2A schematically illustrates the experimental setup used for in vivo retinal imaging of rat eyes. The light source is a visible light LED (e.g., M530L4, Thorlabs Inc., Newton, NJ, USA) with a center wavelength of 530 nm and a full-width half maximum (FWHM) of 35 nm. For albino rats, the contrast of the retinal vasculature is poor when visualized with red and NIR light, and the UV portion of the spectrum is heavily attenuated by the ocular lens. Therefore, the choice of wavelength in the green portion of the spectrum was made as it shows the retinal vessels with excellent contrast. An optical fiber (e.g., MHP550L02, Thorlabs Inc., Newton, NJ, USA) coupled the LED light to the eyelid of the rat and the end of the fiber was fixed to the eyelid with tape. The optical power at the tip of the fiber was measured to be 5 mW.

[0055] The imaging system comprises the transparent ultrasound probe as the frontal element, followed by an ophthalmic lens (L1) (f= 11 mm) (e.g., Volk Digital Series Wide Field lens, Volk, Mentor, OH, USA), a relay lens (L2) (f=-50 mm, e.g., LC1715, Thorlabs Inc., Newton, NJ, USA), and a camera lens (L3) (12 fixed focal length lens, 33- 303, e.g., Edmund Optics Inc., Barrington, NJ). The working distance of L1 was chosen with the thickness of the transparent ultrasound probe in mind, so that, when the transducer touches the cornea properly, a retinal image is created by L1 at the retina conjugate plane (RCP) and this image is relayed to the camera sensor (e.g., GS3-U3- 41S4M-C, FLIR Integrated Imaging Solutions Inc., Richmond, Canada) by L2 and L3. FIG.2B shows a photographic illustration of the imaging system.

[0056] Furthermore, how the transparency of the ultrasound probe might affect the image quality during human eye imaging was investigated. In order to do that, an eye model (e.g., OEMI-7, Ocular Instruments, Bellevue, Washington, USA) was imaged first with the exemplary imaging system. Then the ultrasound transducer was replaced with a glass window (e.g., WG10530, Thorlabs Inc., Newton, NJ, USA). The glass window is similar in thickness to the ultrasound transducer, but optically neutral. An ultrasound signal was generated and received by a pulser receiver (e.g., Panametrics 5900PR,Docket: 320903-2510 Olympus NDT Inc., Waltham, MA, USA) and a custom-made MATLAB script was created to visualize and perform adequate filtering on the signal. A custom-made LABVIEW interface was created to control image acquisition.

[0057] Animal preparation. The Association of Research in Vision and Ophthalmology's guidelines for the ethical use of animals in ophthalmic and visual science research were followed in all aspects of the experimental protocols and related animal care. The related experimental protocol was approved by the University of Illinois Chicago's (UIC) Animal Care Committee. The study employed a sample of eleven- month-old male Sprague Dawley rats (N = 2). The rats were acquired from Charles River Laboratories (Wilmington, MA) and housed within UIC's Biology Resource Lab. After the rats were anesthetized, the pupil was dilated using a drop of 1% tropicamide ophthalmic solution. A lubricant ocular gel (e.g., GenTeal, Alcon Laboratories Inc., Fort Worth, TX, USA) was applied to the whole anterior eye to keep the eye surface hydrated throughout the experiment. The eye gel also served as a coupling medium between the transparent ultrasound probe and the eye to ensure the propagation of the ultrasound waves. An animal holder was used to keep the rat steady and keep the head of the rat immobilized. The animal holder facilitated five degrees of freedom position controls. Results

[0058] FIGS.3A and 3B illustrate the electrical impedance and phase spectrums of the fabricated transparent transducer. The resonant frequency (^^^) and anti-resonant frequency (^^^) were located as 32, and 41, respectively, which is close to the simulation results. Hence, the effective electromechanical coupling coefficient (^^^^^) of the transducer, describing the conversion efficiency between electrical energy and mechanical energy, is calculated as 0.62 by the following equation (2): ^^మ ^^^^^(2) The electric impedance400 around the resonance frequency of the transducer, approximately closing to match the ideal electrical impedance of 50 Ω. The pulse-echo performance of the transducer is illustrated in FIGS.3C and 3D, which show simulated and measured pulse-echo performance of the transducer respectively. The result shows the central frequency of the transducer as 37 MHz, and the -6 db fractional bandwidth as 30%, which are in good agreement with the simulation results (40 MHz). And based on the equation (1), the ^^^௫^^^can be further estimated as 66 µm.Docket: 320903-2510

[0059] FIGS.4A and 4B illustrates the result from an in vivo experiment in a rat eye. FIG.4A shows the fundus image and FIG.4B shows the corresponding ultrasound pulse-echo image. The acquired pulse echo signal constitutes both positive and negative voltage. However, to ensure similarity with the clinical ocular ultrasound pulse- echo signal, the absolute voltage was illustrated instead of the real voltage. It is evident from FIG.4A that the retinal and choroidal vessels as well as the optic nerve of the rat retina are clearly imaged in the fundus photograph. The ultrasound data shows four distinct peaks corresponding to four anatomical locations inside the eye, namely the cornea (P1), the anterior capsule of the crystalline lens (P2), the posterior capsule of the crystalline lens (P3), and the retina (P4). The axial length of the eye could be approximated from the time interval between the corneal peak and the retinal peak, given what is known about the speed of sound in the eye. The speed of sound in the aqueous humor and vitreous humor is around 1532 m / s, whereas the speed of sound inside the ocular lens is around 1641 m / s. For conservative estimation, e that the speed of sound inside the whole eye was assumed to be 1532 m / s. The axial length of the eye can be found using the following equation, Dൌ^௧∗௩ (3) ଶ where Δ^^ is the time intervalspeed of ultrasound. The axial was to 4.902 mm, to the axial length reported by other studies on a similar strain of rats measured by OCT. The slight mismatch in axial length from the reported value could be attributed to the fact that the lens encompasses more than 60% of the volume of the eye in rats. So, a simple approximation of a constant velocity would not be sufficient for calculating the axial length correctly. However, for the human eye, the lens is small relative to the whole eye, so even a simple approximation could give a nearly accurate result. Furthermore, the velocity of sound in different human eye tissues and their interfaces are well characterized.

[0060] FIG.5A illustrates the optical transmission spectrum of the ultrasound transducer. The optical transmission is above 80% in the blue, green and red portions of the spectrum and goes slightly below 80% at the NIR portion of the spectrum. The effect of the transparent transducer on the imaging performance of the system is illustrated in FIGS.5B and 5C. FIG.5B illustrates an image of the retina of an eye model taken with the proposed prototype. FIG.5C illustrates an image of the retina of the same eye model taken after replacing the transparent transducer with an optically neutral glass window.Docket: 320903-2510

[0061] The glass window was chosen in such a way that it didn't change the magnification of imaging and only the difference in imaging performance would be highlighted. The structures present in the eye model, namely the blood vessels, the optic nerve head, and the pigmented ocular lesions on the superior and inferior retina could be easily identified in both images. The mean RGB intensity of the images taken with the transparent transducer and the glass window are 121 and 135, showing the illumination efficiency of the device is not severely affected by the transparent transducer. The ocular vessel branches present in the model eye are clearly seen in both images. However, the smaller vessels that constitute the vessel branches can be differentiated in the image taken with the glass window, whereas the image taken with the transparent transducer doesn't resolve the smaller vessels in the vessel branches clearly. The quality of the image taken with the transparent transducer should be sufficient to accurately guide the ultrasound signal to ocular lesions, which is the goal of the prototype.

[0062] In summary, the feasibility of using a fundus camera to guide the local position of an ultrasound transducer to detect pulse-echo signals from the posterior segment of the eye has been demonstrated. An exemplary high frequency transparent ultrasound transducer was used to facilitate simultaneous fundus and ultrasound imaging. In vivo testing was performed on rats with the prototype and simultaneous fundus image and ultrasound A-scan were gathered. Furthermore, the optical performance and image quality of the prototype on the human eye were examined with a model eye. In addition to providing objective guidance for ultrasound imaging, the fundus image would also provide valuable information such as the presence of orange pigment or drusen overlying an ocular tumor, which are important features in the diagnosis of melanocytic choroidal tumors.

[0063] For this feasibility study, a high-frequency ultrasound probe with a center frequency of 40 MHz was utilized in order to differentiate the cornea, the ocular lens, and the retina clearly in the rat eye. The rat retina is much thinner than the human retina and needs high-frequency ultrasound signal to differentiate different ocular structures properly. The -6 dB axial resolution of the system was calculated to be 66 µm, which is reasonable for imagining the eye of the rat.

[0064] On the other hand, the attenuation of ultrasound waves significantly increases with the operating frequency, reducing the penetration depth. The smaller axial length of a rat eye (about 7 mm) helps to acquire signals from the retina with reasonable quality. For the human eye, the attenuation at this frequency makes the acquisition of signal from the retina impossible. Therefore, ultrasound measurementsDocket: 320903-2510 from the human eye with this preliminary light-ultrasound prototype were not possible. However, the primary purpose of the initial exemplary device and device components herein was to demonstrate the feasibility of using fundus imaging as the guidance for ocular ultrasound with an in vivo experiment on rat eyes, which is clearly demonstrated herein.

[0065] Ultrasound probes used in ophthalmology clinics generally have a center frequency between 7 MHz-20 MHz to compensate for the attenuation to enhance the penetration depth. Therefore, a transparent ultrasound transducer with a frequency similar to the clinical ophthalmic ultrasound systems can be used in a system herein of human eyes. In fact, exemplary systems herein can be designed that operate at lower frequencies, such as 1 to 30 MHz, while other exemplary systems can be designed to operate at higher frequencies, such as 30 to 50 MHz. Alternatively, an exemplary system can be constructed in accordance with the principles herein that is configured to operate in a wide range of frequencies, such as from 1 to 50 MHz, for example.

[0066] The ultrasound transducer used herein was a 5 mm x 5 mm square-shaped transducer. The limited lateral resolution from the transducer means that the echo signal received by the transducer comes from a larger area inside the eye, which in tum makes the signal noise prone. Notably, the signal coming from the posterior of the crystalline lens (peak P3 in FIG.4B) has a relatively weaker signal strength. This may be attributed to the large beam width and the acute curvature of the posterior section of the rat lens as a portion of the beam reaches the border while the rest is still traveling inside the lens, creating an averaging effect.

[0067] An eye model was used to verify the effect that the transparent ultrasound probe had on the image quality (FIGS.5B-5C). For the actual human eye, factors such as pupil size, and eye movement could affect the result. Therefore, the model eye was a better choice for accurate comparison. To deliver the light inside the model eye, a small hole was drilled 5 mm away from the cornea of the model eye and a small optical fiber was inserted which delivered the illumination inside the eye.

[0068] For in vivo testing in rat eyes, the retina was illuminated through the eyelid, which is contrary to the conventional method of both delivering the illumination and receiving the retinal reflection through the pupil. This method has recently been proposed to acquire widefield retinal images even with limited pupil size in human eyes.

[0069] However, the added complexity of having a separate illumination path would increase the complexity of fundus image-guided ultrasound acquisition in patients. Miniaturized indirect illumination-based fundus imagers have shown promise in acquiring non-mydriatic widefield fundus images without the complexity of conventional pupillary ring illumination or a separate illumination system. A simple solution to theDocket: 320903-2510 pupil limitation without requiring a separate illumination system is provided by miniature indirect illumination, which offers a different approach towards the development of widefield portable fundus cameras. Exemplary systems herein can include a miniature, indirect illumination device and method to construct a compact, portable, widefield fundus image-guided ultrasound probe to accurately extract ultrasound information from ocular lesions in human patients.

[0070] A fundus image-guided ultrasound probe has been demonstrated for accurate measurement of ultrasound pulse-echo signal from the posterior segment of the eye. In vivo imaging of rat eyes demonstrated the feasibility of capturing retinal images and ultrasound measurements simultaneously. The optical performance and the image quality of the prototype were further validated using a model eye to demonstrate potential usability in humans.

[0071] Figures 6 through 8 illustrate various components and outputs generated by exemplary systems constructed in accordance with the principles herein. FIG.6 is a conceptual illustration of the proposed hybrid light-ultrasound imager. The integrated ultrasound A-scan is aligned to the center (cross) of the fundus image. The fundus image (a) can be used to readily guide the ultrasound probe to focal lesions L1 in image (b1) and L2 in image (c1) for ultrasound A-scan measurements (b2 and c2). FIG.7A shows a front view of a two-element ultrasound probe for fundus photography guided A-scan measurement and FIG.7B shows a cross-section view of a linear array ultrasound probe for fundus photography guided B-scan measurement. FIG.8 shows a design cross-section of the proposed transparent ultrasonic transducer.

[0072] FIG.9 is a conceptual illustration of the proposed two-element transducer design to enable light fundus photography guided ultrasound examination up to the ora serrata, i.e., the far end of the retina. Two integrated ultrasound A-scans are aligned to the center (US1, cross) and periphery (US2, cross) of the fundus image (a). The fundus image (a) can be used to readily guide the ultrasound probe to focal lesions in images (b1) and (c1) for ultrasound A- scan measurements (b2 and c2). One integrated ultrasound A-scan is aligned to the center (US1, cross) as shown in image (a) of FIG.9. In principle, the fundus camera can be rotated to guide the US1 probe to focal lesions L1 & L2 (images (b1) and (c1) in FIG.9) for ultrasound A-scan measurements (scans (b2) and (c2) in FIG.9). However, for a lesion in peripheral region, such as focal lesion L3 in the image (c1) in FIG.9, a large rotation angle of the imager may produce image distortion and vignetting effect. A second ultrasound A-scan can be aligned to the periphery (US2, cross) of the fundus image (a) of FIG.9 to reach focal lesions in peripheral regions, up to the ora serrata, i.e., the far end of the fundus (see image (d) of FIG.9). Although exemplary embodiments are set forth herein, other embodimentsDocket: 320903-2510 constructed in accordance with the principles of the present disclosure are contemplated as well. Hybrid light-ultrasound fundus imager

[0073] A hybrid light-ultrasound fundus imager was developed and validated for rapid detection and quantitative assessment of ocular lesions. Reliable detection of ocular lesions, such as retinal tears / detachments and ocular tumors, is important for prompt treatment to prevent vision loss. Light fundus photography provides valuable information for eye disease diagnosis and treatment assessment. However, the light- based imaging has limited penetration capability through the ocular tissue. Particularly, ocular media opacity, due to corneal pathology, dense cataract, vitreous hemorrhage, vitritis, or intraocular foreign body from trauma, may limit the light imaging performance for reliable detection of focal lesions. Ultrasound imaging provides a useful supplement for ocular examination. However, the ultrasound resolution and speed are relatively low. The time-consuming ultrasound measurement requires a more skilled ultrasound operator than a typical light fundus photographer to detect and evaluate ocular lesions. Therefore, there is an unmet need for developing a hybrid light-ultrasound fundus imager. The function of the light subsystem is twofold: 1) to work as a traditional fundus camera for clinical assessment of the chorioretinal system; 2) to provide objective guidance for rapid ultrasound measurement of focal lesions of interest in the retina and choroid.

[0074] Miniaturized indirect ophthalmoscopy illumination. Miniaturized indirect illumination can be utilized with the fundus camera. There are two unique differences to differentiate the miniaturized indirect illumination system from the optical design used in conventional fundus cameras. First, instead of using a ring-shape illumination as in conventional fundus cameras, a single-spot of the pupil plane can be used for indirect illumination of the fundus. For the single-spot illumination, roughly half of the pupil size is needed to provide the same buffer range as required in ring-shape illumination to minimize the reflectance artifact. Therefore, the FOV of the fundus camera can be readily enhanced if the same pupil size is available. Second, the indirect ophthalmoscopy illumination can be directly implemented using a miniaturized light source nearby the camera lens. Therefore, the optics in the illumination arm and the hollow mirror in a traditional fundus camera are no longer needed, and the imaging system can be simplified significantly to enable a compact design.

[0075] Contact lens based transparent ultrasound transducer. A contact lens transducer can be used as a kernel part of the hybrid light-ultrasound imager. A contact lens transducer, which is transparent for visible and near-infrared light, can enable easy combination with the light fundus imaging system. The transducer can be opticallyDocket: 320903-2510 designed for ocular imaging, by considering the frequency bandwidth, axial resolution, and penetration capability in the ocular tissue. For example, the transparent ultrasound transducer can be designed with a frequency bandwidth of 17 MHz, which corresponds to about 100 μm axial resolution and about 25 mm penetration capability in the ocular tissue. In various embodiments, the contact lens based transparent transducer can have a central frequency in a range from 15-20 MHz, excellent transparency of up to 90% in the visible wavelength range, an axial resolution of 90-100 µm, a lateral resolution of 80- 100 µm, a high sensitivity of received peak-to-peak output voltage greater than 0.6 Vpp, and easy integration into the fundus camera.

[0076] Handheld portable fundus camera. A portable fundus camera can be used as an imaging platform to add ultrasound capability. Miniaturized indirect ophthalmoscopy illumination can be used for the proposed portable device. The same contact lens, for fabricating the transparent transducer, can be used in the optical design of the fundus camera. The snapshot field of view (FOV) will be designed at 30° visual angle. Flexible rotation of the portable device can readily reach the eyeball equator needed for ultrasound guidance. The ultrasound probe can be aligned to the center (see cross as shown in image (a) of FIG.6) of the FOV of the fundus camera.

[0077] Transparent ultrasound transducers have attracted lots of attentions in various fields, especially in photoacoustic imaging areas, due to its excellent transparency and ultrasound sensitivity. A transparent transducer has been demonstrated and validated for photoacoustic microscopy (PAM). FIG.10A is a cross- section schematically illustrating an example of a transparent ultrasonic transducer and FIG.10B is an image of the transparent ultrasonic transducer. As shown in FIG.10A, the transducer includes a transparent high-frequency (37 MHz) ultrasound transducer with ITO-coated lithium niobate single-crystal. The optical transmission efficiency of the fabricated transparent transducer was evaluated. An optical parametric oscillator laser (e.g., NT242, EKSPLA, Vilnius, Lithuania) was applied to provide the illumination at a broad spectrum, and a laser power meter (e.g., Vega, Ophir Optronics Solutions Ltd., Jerusalem, Israel) was used to measure the ratio of the laser light transmitted through the transparent transducer. In FIG.10C, transmission efficiencies were greater than 80% from 450 to 1064 nm, showing an excellent transparency of the fabricated ultrasound transducer. To better study the performance of the transparent transducer, the image of a resolution target phantom composed of carbon nanotube patterns was achieved. The axial resolution was approximately calculated as 105 µm due to narrow bandwidth. Moreover, an in vivo test was also implemented by imaging a mouse ear to demonstrate the excellent performance of the fabricated transducer. Vessels with a diameter of 30 µm were clearly imaged and displayed.Docket: 320903-2510

[0078] Design of contact lens based transducer. While the transducer in FIG. 10A demonstrated great light transparency, its direct application to the proposed fundus imager was not practical due to the limited physical dimension allowed for ocular imaging. A new high-frequency contact lens based transducer in transparent form is proposed to validate the feasibility of fitting with a fundus camera. FIG.11 is a cross- section schematically illustrating an example of the contact lens based transparent ultrasonic transducer. As shown in FIG.11, a contact lens will work as the substrate for the transducer. A lithium niobate single crystal coated with ITO electrodes can be applied as the core component of the transducer, due to its high optical transmission. Optically transparent epoxy (e.g., EPO-TEK 301, Epoxy Technology, Inc., Billerica, MA, USA) is chosen as the backing layer, between the contact lens and IOT electrode. In the example of FIG.11, a SubMiniature version A (SMA) connector is connected with soft coaxial double shield cable for better electric shielding, while the signal wire and the copper braid on the other side of the cable are connected with lithium niobate wafer and metal housing for signal and ground connection. This design can increase the available space for ocular imaging. The electrical impedance of the contact lens based transducer with a circular type of active piezoelectric material can be designed as 50 ohms to correspond to minimum loss, maximum power, and maximum output voltage of the designed transducer. The transducer has a spherical shape to match to the contact lens surface, and thus to allow easy integration of the ultrasound probe with the fundus camera (see FIGS.15A-15B).

[0079] A high-frequency transparent contact lens transducer was designed and simulated by the KLM transducer equivalent circuit model-based modeling software. The designed parameters of the ultrasound transducer are listed in Table 2. Table 2. Design parameters of the lithium niobate contact lens based transducer in transparent form. Specifications ValuesSimulated pulse-echo performance with bandwidth, electrical impedance magnitude and phase spectrum of the transducer are shown in FIGS.12A and 12B. FIG.12A shows the PiezoCAD simulated pulse-echo and FIG.12B shows the PiezoCAD simulated electrical impedance magnitude and phase spectrum of the designed contact lensDocket: 320903-2510 based transducer. The resonant frequency (^^^) and anti-resonant frequency (^^^) can be located as 14.5 MHz and 19 MHz, respectively. Thus, the effective electromechanical coupling coefficient (^^^^^) of the contact lens based transducer, describing the conversion efficiency between electrical energy and mechanical energy, is calculated by equation (2). The ^^^^^was calculated as 0.65. The electric impedance magnitude of the newly designed transducer is in the range of 41 to 64 Ω around the resonance frequency of the transducer, approximately closing to match the ideal electrical impedance of 50 Ω. The simulated results also show the -6 dB fractional bandwidth as 42%. Theoretically, the -6 dB axial resolution (^^^௫^^^) can be estimated as follows equation (1), where ^^ is the wavelength of the ultrasound in water, and ^^^^ is the -6 dB bandwidth of the designed transducer. Hence, the -6 dB axial resolution can be estimated around 100 μm. With expected high resolution, the contact lens based transducer will better fit the fundus camera for ocular imaging.

[0080] A fabricated transducer will have a central frequency at 17 MHz, -6 dB fractional bandwidth of 42%, ideal electric impedance of 50 ohms, and excellent pulse- echo output voltage of 0.6 Vpp without gain from pulse-echo machine. Hence, the transparent ultrasound transducer offers excellent transparency, axial resolution with 90-100 um of imaging, higher sensitivity.

[0081] The transparent transducer of FIG.11 allows easy integration with the optical imaging system for the proposed hybrid light-ultrasound system. However, the fractional bandwidth of transparent lithium niobate single crystal might be limited due to limitation of the backing materials. The fractional bandwidth can be optimized by design of the transducer and epoxy backing layer thicknesses. A transducer in ring-type form can achieve higher optical transmission for ocular imaging such as shown in FIG.13A. To achieve a sufficient bandwidth, which affects the axial resolution of the ultrasound probe, the lithium niobate single crystal transducer with ring-type form can be used with, e.g., a focal distance of 24 mm. A transducer in ring-type form with outer diameter of 13 mm and inner diameter of 6 mm can achieve higher optical transmission for ocular imaging. As illustrated in FIG.13A, the contact lens for the fundus camera can be inserted into the middle part of the transducer with a hollow structure. For this hollow- shape transducer design, solder can be used as backing and silver-epoxy to obtain the large bandwidth.

[0082] The PiezoCAD simulated pulse-echo performance with bandwidth, electrical impedance magnitude and phase spectrum of the transducer in ring-type form is shown in FIGS.13B and 13C. The central frequency and fractional bandwidth of the transducer are shown as 17 MHz, and 70%, separately. The resonant frequency (^^^) and anti-Docket: 320903-2510 resonant frequency (^^^) can be located as 15.4 MHz and 18.7 MHz, respectively. Thus, based on equation (2), the ^^^^^was determined as 0.57. Moreover, based on equation (1), the -6 dB axial resolution (^^^௫^^^) can be estimated as 63 μm, showing a great potential for high-resolution imaging.

[0083] System design. FIG.14A shows the optical layout of the proposed portable fundus camera with miniaturized indirect illumination. In FIG.14A, L1 is a meniscus lens which contacts the eyeball via water-based clear coupling gel; L2 is a positive lens to produce an intermediate retinal image (dashed vertical line RI in FIG.14A); L3 is the camera lens to relay the retinal image RI to the fundus camera sensor (CS). The light source (LS) is within the same plane as the lens L3. The LS-L3 plane is conjugated to the pupil P of the participant’s eye. In this way, the pupil regions used for illumination (LS’) and imaging can be naturally separated from each other, to ensure the indirect ophthalmoscopy illumination according to the Gullstrand-Principle.

[0084] The fiber coupled light source comprises four LEDs as shown in FIG.14B. The broad band 565 nm light can be used for color fundus imaging (as shown in image (c) of FIG.14C), and narrow band 530 nm, 625 nm, and 780 nm LEDs can be used for multi-spectral imaging (images (d1), (d2), and (d3) of FIG.14C). As shown in image (d1), the 530 nm green light image is predominated by retinal vasculature, the 625 nm red light image (d2) shows mixed retinal and choroidal structures, and the 780 nm image (d3) is predominated by choroidal information.

[0085] Prototype construction. FIG.15A illustrates Zemax simulation of the optical imaging system in FIG.14A. FIG.15B shows a cross-section view of the opto- mechanical arrangement corresponding to FIG.15A, illustrating an example of the portable, compact design. For the simulation in FIG.15A, the L1 is a meniscus lens with -20 mm and -9 mm curvature radiuses of the front and back surfaces, respectively. The clear aperture of the L1 is 10 mm. The L2a and L2b are two plano-convex lenses (e.g., LA4306; Thorlabs) which each have a 1-inch clear aperture and a 20 mm focal length. The L3 is a 16 mm focal length camera lens (37181; Edmund optics), with f / 2.5. As shown in FIG.15A, the overall length of the handheld system can be <150 mm, including the optical lenses and camera. Given the 1-inch diameter of the largest optical lenses, the diameter of the handheld part can be readily controlled within 1.5 inches.

[0086] Technical validation. It is known that the vortex vein ampullas are within the equator region. FIG.16 shows an image (a) of a full fundus of a young adult, an image (b) of a model eye with metal markers, and NIR imaging (c1) and (c2) of the fundus beyond the equator. For preliminary testing of the prototype camera, the model eye with the equator region labeled with metal markers was assembled. While theDocket: 320903-2510 location predefined metal markers can be used to verify the effective FOV of the fundus imager, they can also be used as targets for ultrasound testing. As shown in images (c1) and (c2), the use of NIR illumination for choroidal imaging, which allows direct observation of the vein vortex ampullas (arrows in images (c1) and (c2) of FIG.16).

[0087] The ultrasound probe can be aligned to the center (cross, image (a) of FIG. 6) of the FOV of the fundus camera. The eye model in image (b) of FIG.16 verifies the light-ultrasound alignment. As shown, metal markers were set up in locations corresponding to different visual field regions in the fundus of the model eye. Therefore, the predefined parameters (i.e., location and dimension) of these markers can be used for quantitative calibration and assessment of the light-ultrasound fundus imager. Specifications, including image resolution, FOV and contrast of the fundus camera, ultrasound speed, signal-to-noise ratio (SNR), resolution and penetration range can be quantitatively evaluated.

[0088] The color fundus image (c) of FIG.14C can be used for accurate identification of ocular lesions. However, the requirement for continuous white light illumination to provide real time guidance for ultrasound measurement can be challenging. In a traditional fundus camera, a high-quality image is typically captured in a snapshot mode, i.e., delivering a high-powered light flash to capture a color fundus image in a short exposure time window. Unfortunately, the flash illumination mode for snapshot imaging is not feasible for the proposed device, because of the need for real- time image guidance. To tackle this challenge, a 565 nm LED was selected as the light source to minimize the photochemical hazard risk to the retina. By excluding the short wavelength light to the retina, the photochemical hazard risk for light wavelength > 500 nm can be significantly reduced. As shown in FIG.14B, the red illumination shows mixed retinal and choroidal structures, while the NIR light is predominated by the choroidal vasculature. The red and NIR light efficiencies are significantly better than the short blue / green light. This is why color images are typically red oriented in traditional fundus cameras with white light illumination. Moreover, the aphakic photochemical hazard weighting function is quickly reduced with increasing wavelength. Therefore, red or NIR video imaging can be achieved without light safety concern.

[0089] Data acquisition and analysis. Lesion basal diameter was be measured using the conventional widefield light fundus camera using the Ruler measurement tool of the Optos Advanced software. Image (a) of FIG.17 shows a conventional widefield light fundus photo of a choroidal halo nevus with lesion basal diameter measurements (green calipers). Image (b) of FIG.17 shows the conventional ultrasound with lesion thickness measurement (green caliper) and image (c) shows the enhanced depth imaging-OCT with lesion thickness measurement (green caliper). Lesion thickness onDocket: 320903-2510 EDI-OCT can be measured if the lesion is less than 1.5 mm in thickness using the Distance Measure function of the Heidelberg Eye Explorer software. Image (d1) of FIG. 17 shows an example of the proposed hybrid light-ultrasound imager aligned to the center of the choroidal nevus on fundus image (cross) with image (d2) showing an A- scan measurement (green caliper). Lesion thickness can be measured on the conventional ophthalmic ultrasound A-scan using the caliper function of the Quantel Medical software. Internal acoustic impedance can be measured on the conventional ophthalmic ultrasound standardized A-scan.

[0090] Ultrasound A-scan is an important tool for quantitative assessment of ocular lesions. However, its usability is limited by the difficulty of accurately localizing the ultrasound probe to a lesion of interest. In accordance with the principles of the present disclosure, a transparent LiNbO3single crystal ultrasound transducer was fabricated, and integrated with a widefield fundus camera to guide the ultrasound local position. The electrical impedance, phase spectrum, pulse-echo performance, and optical transmission spectrum of the ultrasound transducer were validated. The fundus camera guided ultrasound probe imaging system herein was tested for in vivo measurement of rat eyes. Anterior and posterior segments of the rat eye could be unambiguously differentiated with the fundus photography guided ultrasound measurement. A model eye was also used to verify the imaging performance of the prototype device in the human eye. The prototype shows the potential of being used in the clinic to accurately measure the thickness and echogenicity of ocular lesions in vivo.

[0091] It should be emphasized that the above-described embodiments of the present disclosure are merely possible examples of implementations set forth for a clear understanding of the principles of the disclosure. Many variations and modifications may be made to the above-described embodiment(s) without departing substantially from the spirit and principles of the disclosure. All such modifications and variations are intended to be included herein within the scope of this disclosure and protected by the following claims.

[0092] The term "substantially" is meant to permit deviations from the descriptive term that don't negatively impact the intended purpose. Descriptive terms are implicitly understood to be modified by the word substantially, even if the term is not explicitly modified by the word substantially.

[0093] It should be noted that ratios, concentrations, amounts, and other numerical data may be expressed herein in a range format. It is to be understood that such a range format is used for convenience and brevity, and thus, should be interpreted in a flexible manner to include not only the numerical values explicitly recited as the limits of the range, but also to include all the individual numerical values or sub-rangesDocket: 320903-2510 encompassed within that range as if each numerical value and sub-range is explicitly recited. To illustrate, a concentration range of “about 0.1% to about 5%” should be interpreted to include not only the explicitly recited concentration of about 0.1 wt% to about 5 wt%, but also include individual concentrations (e.g., 1%, 2%, 3%, and 4%) and the sub-ranges (e.g., 0.5%, 1.1%, 2.2%, 3.3%, and 4.4%) within the indicated range. The term “about” can include traditional rounding according to significant figures of numerical values. In addition, the phrase “about ‘x’ to ‘y’” includes “about ‘x’ to about ‘y’”.

Claims

Docket: 320903-2510 CLAIMS Therefore, at least the following is claimed:

1. An ocular imaging system, comprising: a substantially transparent ultrasound transducer; an optical imaging system comprising a fundus camera sensor aligned with the ultrasound transducer; and one or more control interface components configured to enable visualization of an ocular surface and perform filtering on imaging signals generated by the ultrasound transducer operatively connected to the fundus camera sensor.

2. The ocular imaging system of claim 1, wherein the optical imaging system comprises a miniature indirect illumination source operatively connectable to the one or more control interface components.

3. The ocular imaging system of claim 2, wherein the miniature indirect illumination source comprises a source of light positioned to illuminate the ocular surface via one or more lens of the optical imaging system.

4. The ocular imaging system of claim 3, wherein the source of light comprises a combination of broadband and narrowband light emitting diodes.

5. The ocular imaging system of claim 3, wherein the source of light is configured to provide light at wavelengths greater than 500 nm.

6. The ocular imaging system of claim 3, wherein the source of light is in a plane with and offset from a camera lens of the fundus camera sensor.

7. The ocular imaging system of claim 3, wherein the source of light illuminates the ocular surface via the substantially transparent ultrasound transducer.

8. The ocular imaging system of any one of claims 1-7, wherein the substantially transparent ultrasound transducer is a single ultrasound transducer.Docket: 320903-2510 9. The ocular imaging system of claim 8, wherein the substantially transparent ultrasound transducer comprises a contact lens disposed across a side of the single ultrasound transducer adjacent to the fundus camera sensor.

10. The ocular imaging system of claim 9, wherein the single ultrasound transducer is curved.

11. The ocular imaging system of claim 8, wherein the substantially transparent ultrasound transducer comprises a contact lens disposed within a hollow middle portion of the single ultrasound transducer.

12. The ocular imaging system of any one of claims 1-7, comprising a second substantially transparent ultrasound transducer and a contact lens disposed across both ultrasound transducers.

13. The ocular imaging system of any of claims 1-7, wherein the substantially transparent ultrasound transducer comprises a lithium niobate (LNO) crystal with indium tin oxide (ITO) electrodes disposed on opposite sides of the LNO crystal.

14. The ocular imaging system of claim 13, wherein the substantially transparent ultrasound transducer comprises a matching layer disposed on a side opposite the fundus camera sensor.

15. The ocular imaging system of any of claims 1-14, wherein the ocular imaging system is a portable handheld imaging system.

16. A method for ocular imaging, comprising: illuminating an ocular surface of an eye with the ocular imaging system of any of claims 1-15; aligning the ultrasound transducer with a lesion on the ocular surface using the optical imaging system; and obtaining ultrasound pulse-echo scan data of the lesion using the ultrasound transducer.

17. The method of claim 16, comprising obtaining ultrasound pulse-echo scan data at a periphery of an ocular image using a second substantially transparent ultrasound transducer.Docket: 320903-2510 18. The method of claim 16, wherein the ocular surface is illuminated with light at wavelengths greater than 500 nm.

19. The method of claim 16, wherein an illumination efficiency is configurable to accurately guide the ultrasound transducer to the lesion on the ocular surface located in a posterior region of the eye.

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