Abstract:
A system and method for imaging retinal tissues in an eye generates an input light beam having ultra-short pulses and an input wavelength (λ i ) to stimulate the tissue. Depending on the particular type tissue being imaged, the retinal tissue responds to the input beam by generating a return beam of light having first and second components of different wavelengths (λ r1  and λ r2 ). An imaging unit then receives the return light and images the tissue according to the return wavelength (λ r1  vis-a-vis λ r2 ). Additionally, a sensor unit is used to evaluate light returning from the retina to measure optical and phase aberrations introduced by the eye, and to program a compensator (e.g. an active mirror) that compensates the input beam by removing the aberrations.

Description:
FIELD OF THE INVENTION 
   The present invention pertains generally to ophthalmic diagnostic equipment. More particularly, the present invention pertains to systems and methods for imaging retinal tissue. The present invention is particularly, but not exclusively, useful as a system and method for stimulating tissue with a light beam of ultra-short pulses having an input wavelength that generates a return beam having different wavelength components depending on the type of retinal tissue being imaged. 
   BACKGROUND OF THE INVENTION 
   Effective imaging of the retina of an eye depends on the type of retinal tissue that is to be imaged, as well as the optical response of that tissue to the input light beam. In particular, for two specific tissues of the retina, namely the Retina Pigment Epithelium (RPE) and the Lamina Cribrosa (LC), it happens there are two different optical phenomena that generate the particular tissue&#39;s response. One is known as Two Photon Excited Fluorescence (TPEF). This phenomenon is efficacious for imaging the RPE of the retina. The other phenomenon is Second Harmonic Generation (SHG), which is efficacious for imaging the LC. An ability to image these tissues (i.e. RPE or LC) depends on how these phenomena are exploited. 
   Anatomically, RPE tissue in the retina includes the protein, lipofuscin. In the context of the present invention, it is known that lipofuscin is susceptible to TPEF. Specifically, it can be demonstrated that when an input beam of red light (e.g. λ i =780 nm) is incident on lipofuscin in the RPE, a resultant return beam of fluorescent green light (e.g. λ r1 =530 nm) is generated. On the other hand, when this same input beam of red light (λ i ) is incident on the LC there is a much different response. Specifically, as a result of SHG, a return beam of blue light (e.g. % λ r2 =390 nm) is generated. (Note: λ i ≠λ r1 ≠λ r2 ). Nevertheless, both of the return beams (λ r1  and λ r2 ) are useable for effectively imaging the respective tissues. 
   During an imaging procedure, it happens that the anterior components of the eye (i.e. the cornea and the lens) will introduce optical aberrations into the input light beam. Also, the retina will introduce optical and phase aberrations. These aberrations, both optical and phase aberrations, are measurable. Furthermore, using adaptive optics with a wavefront sensor, the input light can be altered to effectively compensate for any optical aberrations that may be present. Further, phase aberrations that are introduced by curvature of the retina can be compensated for by pre-programming input to a computer that controls the adaptive optics. 
   In light of the above, it is an object of the present invention to provide a system and method that is capable of alternatively imaging the RPE or the LC tissues in a retina of an eye. Another object of the present invention is to provide a system and method that is capable of selectively exploiting the TPEF or SHG phenomenon to image different tissue in the retina of an eye. Yet another object of the present invention is to provide a system and method for imaging selective tissue in the retina of an eye that is easy to implement, is simple to use and is comparatively cost effective. 
   SUMMARY OF THE INVENTION 
   A system and method for imaging tissue in the retina of an eye includes a laser unit for generating an ultra-short pulsed input light beam having a wavelength (λ i ). As envisioned for the present invention, when the input light beam (λ i ) is incident on a target tissue, the tissue will generate a return light beam (λ r ). Importantly, this return light beam will include different wavelength components (i.e. λ r1  and λ r2 ) depending on the nature of the target tissue. In accordance with the present invention, this return light beam is then used for two different purposes. For one, regardless of wavelength, the return beam includes information that can be used to compensate for optical and phase aberrations that are introduced into the input beam by the eye. For another, depending on which component of the return beam is predominant (i.e. λ r1  vis-a-vis λ r2 ) the selected component can be used to image the particular retinal tissue that generates the return light beam. 
   Structurally, along with the laser unit that is used for generating the input light beam, the system for the present invention also includes a sensor with adaptive optics. For the present invention, the sensor has a wavefront sensor for measuring optical aberrations (e.g. a Hartmann Shack sensor) that is electronically connected with an active mirror. Together, the wavefront sensor and the active mirror are employed to alter the input light beam in a manner that will compensate for optical and phase aberrations introduced into the input beam. The system also includes a detector that receives the return light beam and uses it for imaging the target tissue that has been illuminated by the input beam. 
   For imaging purposes, the present invention directs the input light beam onto the target tissue that is to be imaged (e.g. RPE or LC). Preferably the input light beam is red light having a wavelength of about λ i =780 nm. In the case of the RPE, because the target tissue includes lipofuscin, the tissue responds with TPEF by generating a return beam of green fluorescent light (λ r1 =580 nm). In the case of the LC, however, the target tissue responds with SHG by generating a return beam of blue light (λ r2 =390 nm). In each case, regardless of the type tissue being imaged, the return light is received by the detector for subsequent imaging of the tissue. 

   
     BRIEF DESCRIPTION OF THE DRAWINGS 
     The novel features of this invention, as well as the invention itself, both as to its structure and its operation, will be best understood from the accompanying drawings, taken in conjunction with the accompanying description, in which similar reference characters refer to similar parts, and in which: 
       FIG. 1  is a schematic view of the components of a system for the present invention; 
       FIG. 2  is a cross sectional view of a portion of a retina of an eye; 
       FIG. 3  is an enlarged view of retinal tissue (i.e. RPE) in the area bounded by the line  3 - 3  in  FIG. 2 ; 
       FIG. 4  is an enlarged view of retinal tissue (i.e. LC) in the optical nerve head; and 
       FIG. 5  is a schematic of the aberration compensation mechanism of the system. 
   

   DESCRIPTION OF THE PREFERRED EMBODIMENTS 
   Referring initially to  FIG. 1 , a system in accordance with the present invention is shown and is generally designated  10 . More specifically, as shown, the system  10  includes a laser unit  12  for generating an input laser beam  14 . For the present invention, the input laser beam  14  is preferably a pulsed laser beam wherein the pulses are ultra-short and each pulse has a duration measured in femto-seconds. Further, the input laser beam  14  preferably has a wavelength (λ i ) that is about 780 nm (λ i =780 nm).  FIG. 1  also indicates that the input laser beam  14  is directed from the laser unit  12 , and onto the retina  16  of an eye  18 . As intended for the present invention, when the input light beam  14  (λ i ) is incident on tissue in the retina  16 , it will interact with the tissue to generate a return light beam  20 . Importantly, the return light beam  20  may include either, or both, of two different components that will have different wavelengths. Stated differently, the return light beam  20  will include a first component having a wavelength (λ r1 ) and a second component with a wavelength (λ r2 ). Note: λ i ≠λ r1 ≠λ r2 . 
   Still referring to  FIG. 1  it will be seen that, in addition to the laser unit  12 , the system  10  includes a sensor unit  22  and an active mirror  24 . Specifically, these elements of the system  10  (i.e. sensor unit  22  and active mirror  24 ) are used to pre-compensate the input beam  14  to create a diffraction limited spot on the retina  16 . On this point it is well known that the cornea  26  and lens  28  of the eye  18  will introduce optical aberrations into the input light beam  14 . Also, the retina  16  will introduce phase aberrations that continue with the return light beam  20 . In order to measure the optical aberrations, the sensor unit  22  is preferably a wavefront sensor of a type well known in the pertinent art, such as a Hartmann Shack sensor. On the other hand, phase aberrations introduced by the retina  16  are preferably compensated for by pre-programming a computer to account for curvature of the retina  16 . It is known, however, that some phase aberrations can be detected by fluorescence wavefront analysis. Therefore, the sensor unit  22  may also include this capability. 
   Once optical and phase aberrations in a return light beam  20  have been measured by the sensor unit  22 , the aberrations can then be used to program an active mirror  24  (i.e. the computer used for operation of the active mirror  24 ). Specifically, the active mirror  24  is to be programmed in a manner that will change the input light beam  14  to thereby effectively remove the aberrations from the return light beam  20 . Alternatively, a customized phase plate  29  (see  FIG. 5 ) of a type disclosed in co-pending U.S. application Ser. No. 12/204,674 which is assigned to the same assignee as the present invention can be used with, or without, the active mirror  24  for this purpose. Importantly, the now-compensated return light beam  20  can be used by the imaging unit  30  for imaging purposes. 
   Anatomically, an optic (visual) part  32  of the retina  16  comprises most of what is generally referred to as the fundus. As shown in  FIG. 2 , the sclera  34  is under the optic (visual) part  32 , and the optical nerve head  36  connects to the optic (visual) part  32  through the sclera  34 . In detail, with reference to  FIG. 2  and  FIG. 3  it will be seen that the optic (visual) part  32  of the retina  16  is curved and includes a Retina-Pigment-Epithelium (RPE)  38 . The RPE  38  is a target tissue of interest for the present invention. Anterior to the RPE  38  and identified in an anterior to posterior direction, are: nerve fibers  40 ; retinal ganglion cells  42 ; axion  44 ; bipolar cell  46 ; and a photo receptor  48 . Of these, as indicated above, it is the RPE  38  with its lipofuscins that responds to the input beam (λ i ) to generate a return beam (λ r1 )  20  due to TPEF. Referring now to  FIG. 4 , it will be seen that the optical nerve head  36  anatomically includes the Lamina Cribrosa (LC)  50  which is bounded by pre-laminar tissue  52  and post-laminar tissue  54 . As also indicated above, the LC  50  is also a target tissue of interest for the present invention. In this case, the LC  50  responds to the input beam  14  (λ i ) to generate a return beam  20  (λ r2 ) due to SHG. 
   Additional aspects of aberration compensation for the present invention can be appreciated with reference to  FIG. 5 . There the sensor unit  22  is shown to include a lens array  56 , and a CCD camera  58 . This arrangement is typical for a wavefront sensor of the type commonly referred to as a Hartmann-Shack sensor.  FIG. 5  also indicates that a customized phase plate  29  can be used together with, or in lieu of, the active mirror  24 . In either case, the importance of the arrangement is to compensate the input beam  14  for aberrations that could otherwise diminish the efficacy of the imaging system  10 . Anatomically, there are three sources for these aberrations; all from the eye  18  itself. They are: 1] optical aberrations introduced by the anterior segment (i.e. cornea  26  and lens  28 ); 2] phase aberrations introduced by the curvature of the retina  16  that relate to astigmatism; and 3] phase aberrations introduced by the retina  16 . 
   Of all the aberrations introduced by an eye  18  into the input light beam  14 , optical aberrations are the most prominent, and are measured by the sensor unit  22 . To do this, a source  60  of infrared (IR) light radiates IR through pupil imaging optics  62 . Also, the Internal Limiting Membrane (ILM)  64  that defines the anterior surface of the retina  16  includes aberrational information in the light that is reflected from the retina  16 . After leaving the eye  18 , the optical aberrations that are introduced into the return beam  20  by the cornea  26  and lens  28  are processed by the sensor unit  22 . The resultant information is then programmed into the active mirror  24 . This essentially compensates for the first source of aberrations (i.e. the anterior segments). As for the second source of aberrations (i.e. phase aberrations introduced by the curvature of the retina  16 ) it is well known that these aberrations can be measured in accordance with the angle of incidence, “θ”, between the input light beam  14  and the anterior surface of the retina  16 . Accordingly, “θ” is determined by anatomical dimensions of the retina  16 . The resultant measurements involving “θ” are then also programmed into the computer-controlled active mirror  24 . The remaining aberrations from the third source (i.e. the retina  16 ), although relatively minor, can be detected by a fluorescence wavefront sensor in the sensor unit  22  and used with the other information to further refine compensation corrections for the system  10 . 
   As mentioned above, and as shown in  FIG. 5 , a custom phase plate  29  can be used in combination with the active mirror  24 , or in lieu thereof. In either configuration, the purpose is to pre-compensate the input light beam  14  so that aberrations introduced into the light beam  14  do not detract from the imaging capability of the system  10 . 
   While the particular System and Method for Imaging Retinal Tissue with Tissue Generated Light as herein shown and disclosed in detail is fully capable of obtaining the objects and providing the advantages herein before stated, it is to be understood that it is merely illustrative of the presently preferred embodiments of the invention and that no limitations are intended to the details of construction or design herein shown other than as described in the appended claims.