|Publication number||WO2000071991 A1|
|Publication date||30 Nov 2000|
|Filing date||19 May 2000|
|Priority date||25 May 1999|
|Publication number||PCT/2000/14033, PCT/US/0/014033, PCT/US/0/14033, PCT/US/2000/014033, PCT/US/2000/14033, PCT/US0/014033, PCT/US0/14033, PCT/US0014033, PCT/US014033, PCT/US2000/014033, PCT/US2000/14033, PCT/US2000014033, PCT/US200014033, WO 0071991 A1, WO 0071991A1, WO 2000/071991 A1, WO 2000071991 A1, WO 2000071991A1, WO-A1-0071991, WO-A1-2000071991, WO0071991 A1, WO0071991A1, WO2000/071991A1, WO2000071991 A1, WO2000071991A1|
|Inventors||Thomas M. Baer, Louis J. Dietz, Robert S. Dubrow, Paul G. Hayter, Michael Hodges, Bala S. Manian, Robert J. Shartle, David M. Heffelfinger|
|Applicant||Biometric Imaging, Inc.|
|Export Citation||BiBTeX, EndNote, RefMan|
|Patent Citations (6), Non-Patent Citations (2), Referenced by (29), Classifications (11), Legal Events (6)|
|External Links: Patentscope, Espacenet|
APPARATUS AND METHOD FOR OPTICAL DETECTION IN A LIMITED DEPTH OF FIELD
This invention relates to optical detection instruments more particularly optical detection systems which detect fluorescence in a limited depth of field.
Rapid identification and enumeration of the various components of biological fluids is an important research and diagnostic aim. Minimal processing and handling of samples would contribute to the widespread use of such techniques .
In the case of enumeration of leukocyte subclasses of human blood, the need for improved techniques is especially keen. For example, the usefulness of moni- toring CD4+ lymphocyte levels in noting the progression from HIV positive status to AIDS has underscored the need for a fast, inexpensive, and reliable method to analyze patient blood samples.
Landay et al . , "Application of flow cytometry to the study of HIV infection," AIDS 4:479-497 (1990) describes the utility of a technique in understanding the biology of HIV infection. Multiple-color flow cytometric analysis can be applied to the study of HIV disease by using various monoclonal antibodies to perform phenotypic analysis of blood samples. This technique is also useful in other immune system determinations, as in evaluating the status of organ transplant or leukemia patients.
Flow cytometry is a well-known technique wherein cells may be characterized and separated based on fluorescent emission. A labeled, mono-dispersed cell suspension travels through a tube in a fine fluid stream and is presented to an excitation beam. The emitted fluorescence of each cell is measured by appropriate detectors and the cells may be split into droplets and sorted according to given parameters by electrical and mechanical means .
Flow cytometry may be used to identify and enumerate specific subclasses of blood cells. For example, in U.S. Pat. No. 4,284,412 Hansen et al . , lymphocytes which have been reacted with fluorescently-labeled monoclonal antibodies are separated from red blood cells and presented one by one to a fixed detector in a flow cytometry system. Each cell is characterized by analysis of forward light scatter, right angle scatter, and fluorescence. This method requires complex sample preparation and instrumentation. While flow cytometry has improved assay reliability and reproducibility in this application, it generally cannot directly provide absolute cell counts for lymphocyte subsets. Independent white blood counts and differential white counts are required to calculate absolute cell counts per unit volume. In the usual flow cytometry practice, in order to distinguish lymphocytes from monocytes and granulocytes, a lymphocyte gate based on forward and side light scatter patterns must be established for each sample.
Flow cytometry is not routinely used for identifying and enumerating lymphocyte subclasses in the presence of red blood cells, although U.S. Patent No. 4,727,020 Recktenwald provides a contrary example. Removal of the red blood cells, by density-gradient separation or lysing, increases the time, cost and number of blood-handling steps per assay. Additional blood-han- dling steps increase the potential for exposure to blood-borne infectious agents. As stated above, the resultant data produced by the flow cytometry method is inadequate for some purposes. In order to calculate absolute cell count per unit volume, flow cytometric data must generally be combined with additional data obtained from other methods. Also, because flow cytometers conventionally utilize a fluid stream passing through a small nozzle, they may generate aerosols which pose an additional source of biohazardous materials for laboratory personnel .
An alternative is to fix sample position relative to the excitation beam. For example, in U.S. Pat. No. 4,758,727 and its divisional, U.S. Pat. No. 4,877,966, Tomei et al . , a method and apparatus for measurement of low-level laser-induced fluorescence is described. In this invention, a coherent laser beam is passed through a three-dimensional scanner and focused onto a static target. The target is an object such as a monolayer cell culture or tissue section. A beam spot, having a size as small as one micron, is passed back and forth across the target by a scanner whose path and movement rate are computer-controlled. Fluorescent light is gathered by a biased-cut fiberoptic base plate and relayed to a detector positioned on the opposite side of the target from the beam.
U.S. Pat. No. 5,037,207, also granted to Tomei et al . , discloses a laser imaging system with enhanced spatial resolution and light gathering efficiency which allows for digital imaging of a target of varying size, dependent upon the data retrieval and storage limitations of the supporting computer system. The system utilizes a novel optical fiber detector assembly and a rapid scan for collection of all light from every laser spot to create a quantitative digital reproduction of the image on the surface of a target .
U.S. Pat. Nos. 5,072,382, Kamentsky, and 5,107,422, Kamentsky et al . , disclose an apparatus and method for scanning a cell population with a beam to generate multiparameter optical data based on each cell's specific location. The scan is made of a surface on which cells have been deposited. A background level is estimated for the neighborhood surrounding each cell based on digital data and corrections are made for the background level .
In "Acousto-Optic Laser-Scanning Cytometer, " Cytometry 9:101-110 (1988) Burger and Gershman and U.S. Pat. No. 4,665,553 Gershman et al . , a laser-scanning cytometer is disclosed. An optical scan is made of a lysed and washed sample in a cuvette by a Bragg cell-controlled scanner. The cuvette is translated in a stepwise fashion in one direction relative to the scanner. The scanner operates in a direction perpendicular to the direction of cuvette translation and the scan occurs along the side of the cuvette. Once a cell is located, a beam optimization algorithm operates to steady the beam on the cell and measurements of forward light scatter, orthogo- nal light scatter, and fluorescence are made. Then the process is repeated.
In U.S. Pat. No. 5,117,466, Buican et al . describe a fluorescence analysis system in which data from a flow cytometer establish identification criteria used by a confocal laser microscope to virtually sort the cellular components of a sample. Birefringent optics and Fourier-Transform technology are used to visually select and display cells or subcellular structures having the desired spectral properties.
In "Fluorescence Analysis of Picoliter Samples," Analytical Biochemistry 102:90-96 (1980) Mroz and Lechene teach a method of handling picoliter-volume samples to gather fluorescence intensity data. Samples are taken up via syringe in a single siliconized capillary tube with oil between the samples. Measurements are made of an optical fluorescence chamber defined by a pinhole diaphragm, a microscope objective, and the diameter of the capillary tube.
U.S. patents granted to Mathies et al . are also relevant to the field of the present invention. In U.S. Pat. No. 4,979,824, a high sensitivity detection appara- tus is described. This apparatus is based on a flow cytometry system and utilizes a spatial filter to define a small probe volume that allows for detection of individual fluorescent particles and molecules. Laser power and exposure time of the sample are chosen for the best signal-to-noise ratio. Real-time detection of photon bursts from fluorescent particles is used to distinguish the number, location or concentration of the particles from background energy.
In U.S. Pat. No. 5,091,652 Mathies et al . , a laser-excited fluorescent scanner is revealed for scanning separated samples using a confocal microscope. The sample is preferably separated by and detected from an electrophoresed slab gel, but may also be on a membrane, filter paper, petri dish, or glass substrate. The confo- cal microscope forms an illumination volume in the gel and the beam is oriented so that background scattering is minimized by the polarization characteristics of the scattered light. U.S. Pat. No. 5,274,240 also granted to Mathies et al . and a continuation-in-part of the above patent, teaches a laser-excited capillary array scanner. This invention is primarily intended for fluorescence detec- tion from an array of capillary tubes containing samples that have been separated by capillary electrophoresis . The fluorescence detection assembly employs a confocal system to detect fluorescence from the interior volumes of each capillary tube. The current cytometry art generally requires time-consuming and potentially hazardous sample-handling and component separation steps. It fails to allow for rapid volumetric identification and enumeration of sub- populations of a cell suspension that are present within a mixed population. The techniques of the prior art often require trained personnel .
It is therefore an object of the present invention to provide a quick, simple to use, less expensive, safer, automated apparatus and method for directly ob- taining fluorescence measurements in a limited field of depth adaptable to recognition of fluorescence characteristic of cells, cell fragments or other targets of discrete size. The measurements should be made in fluids and should be adaptable to being made in a volumetric manner enabling assays which require small volumes of sample and reagent.
SUMMARY OF THE INVENTION
The above object has been achieved with an apparatus and method for scanning of a fluid in a limited depth of field which can detect formation of fluorescent complexes. The apparatus is adapted for the optical scanning of a sample holder containing a sample in a static and minimally processed form. The fluorescence is detected from throughout a limited depth of field of a non-flowing sample suspension. Enumeration may be done in a precise volume for the purpose of obtaining absolute cell counts. "Absolute", as defined herein, means the absolute number of cells or other discrete targets such as beads per volume as represented by the volume scanned. As defined herein, "cell" means a whole cell or a part of a cell (i.e. fragments, membranes, organelles, etc.). The complexes are the result of a reaction between fluorescently-labeled binding agents and corresponding binding sites present in cellular components, present on a solid support, or in fluid. An excitation laser beam is directed by an optical scanner to a columnar region in the sample container. In one embodiment, the columnar region is generally defined within the interior depth dimension of a capillary tube by the beam spot of the laser and a spatial filter of sufficient pinhole aperture. The aperture is chosen to selectively detect the fluorescence emitted throughout the columnar region illuminated and is disposed between the capillary tube and a detection means. Because no separation of bound and unbound fluorescently-labeled binding agent is necessary in the sample, both are viewed as fluorescence by the detection means. However, areas of heightened fluorescence intensity occur where the labeled binding agents congregate, namely on the binding sites present in the cellular components of the sample. The detection means, therefore, records a signal of heightened fluorescence intensity above a given threshold of background fluorescence as corresponding to a single cell or other discrete target. In one embodiment, a laser creates an excitation beam of a wavelength of 400 to 1000 nanometers and is focused onto a capillary tube of rectangular cross-section from a position directly above the capil- lary tube. The spot size of the laser beam at the point of its intersection with the capillary tube is 5 to 15 microns in diameter, depending upon the expected cell size, and the illuminated depth dimension in the capillary tube is 25 to 225 microns. As described later, there is a relationship between the spot size and the illuminated depth dimension. In the present invention, an excitation beam is continuously scanned in two directions to impinge upon the outer wall of a transparent capillary tube. The first scan direction follows a path transverse to the longitudinal axis of the capillary tube, i.e. the width of the rectangular cross-section of the capillary tube, and begins and ends at points that are beyond the lateral boundaries of the capillary tube. The second scan direction follows a path along the longitudinal axis of the capillary tube, i.e. a length of the capillary. The scan of a known volume of the capillary tube, achieved by measuring the beginning and ending points in the second scan direction or by beginning and ending the scan at defined points, can be used to calculate the presence of a particular subpopulation of cellular components or other discrete fluorescent targets per unit volume.
The apparatus of the present invention is especially well-suited to the detection of subclasses of blood cells. In a typical assay, a sample of whole uncoagulated blood is obtained and incubated with an excess amount of fluorescently-labeled antibodies that are directed toward various cell surface markers present on blood cell subclasses. The fluorophores are chosen so that they will activate in the wavelength range of the excitation beam. This wavelength range has also been specifically selected to minimize interference due to autofluorescence from blood components not of interest. The sample containing fluorescently-labeled antibody in both complexed and free form is generally diluted and then inserted into the capillary tube. The tube is then optically scanned at wavelengths necessary to excite the fluorophores. Based on fluorescent emission from specific fluorophores used to label specific antibodies, the number of cells of a certain type per unit volume can be quickly determined, as can ratios of cell types present in the blood or other biological fluid sample. The instrument and technique of the present invention quickly detect cellular components of biological fluids in a precise volume and require minimal processing of the sample. The present invention substantially cuts down on assay times and costs and requires mini- mal handling of samples, an especially important precaution during the examination of blood samples. Because no special instruments are necessary for processing the samples and the number of requisite reagents is kept to a minimum, the present invention is well-adapted for use in a clinical setting.
BRIEF DESCRIPTION OF THE DRAWINGS
Fig. 1 is a plan view of an apparatus according to the present invention. Fig. 1A shows a plan view of a fiber optic link used to provide the excitation beam of Fig. 1.
Fig. IB shows a plan view of an acousto-optic scan assembly. Fig. 1C shows a plan view of a dispersive prism used with a segmented detector as the detection means of Fig. 1.
Fig. ID shows a plan view of a direct vision prism used with a segmented detector as the detection means of Fig. 1.
Fig. 2 is a side view of a sample-filled capillary tube illustrating an illuminated columnar region, and both excitation and emission beams. Fig. 2A illustrates the Gaussian waist, limited depth of focus, and relationship of spot size to detection depth in a container as shown in Fig. 2.
Fig. 2B is the view of Fig. 2 with alternate spatial filter aperture size. Fig. 2C illustrates the effect of lens selection on focal length, beam spot size and illuminated depth .
Fig. 3 is a perspective view of a sample-filled capillary tube showing overlapping beam spots and an il- luminated columnar region according to the present invention.
Fig. 4 is a top view of a sample-filled capillary tube, showing overlapping detection of columnar areas according to the present invention. Fig. 5 is a schematic representation of the optical scanning path according to a preferred embodiment of the present invention.
Fig. 6 is a schematic representation of a labeled cell suspension and the corresponding detector signal.
Fig. 7 is a cut-away view of a fiber optic connector shown in plan view in Fig. 1A. BEST MODE FOR CARRYING OUT THE INVENTION
With reference to Fig. 1, laser 10 generates an excitation beam 80. As shown in Fig. 1, a single laser source may be used to produce a single characteristic beam of coherent light of a specific wavelength. At times it may be optimal to substitute light sources or to use multiple light sources for the excitation beam. Fig. 1A provides a plan view for a fiber optic link which allows multiple light sources to be used. In Fig. 1A, light is provided by lasers 271,
272, 273. Each laser can produce light at different wavelengths. The lasers are arranged such that laser 271 has the highest wavelength, laser 272 has an intermediate wavelength and laser 273 has the shortest wavelength. Lasers 271, 272, 273 focus light into optical fibers 203, 202, 201 respectively. Optical fibers 203, 202, 201 each terminate in a respective fiber connector output 214, 212, 210. Each of fiber connector outputs 214, 212, 210 may be removably coupled to fiber connector inputs 215, 213, 211. Optical fibers 219, 218, 217 extend into the optical scanner. It is apparent from this arrangement that any of lasers 271, 272, 273 may be coupled to the optical scanner allowing each laser to be used either alone or in combination. In addition, exchange of lasers is greatly simplified. The fiber connector output can be easily decoupled from the fiber connector input to allow for simplified exchange of lasers.
The preferred wavelengths for use with the optical scanner may be in the 400 nm - 1000 nm range. The optical fibers used with each laser may be adapted to have single-mode light propagation with the guided mode selected to correspond to the light wavelength produced by the corresponding laser. The fiber connector input and fiber output connectors are selected to allow removable coupling of optical fibers into the system. A connector that encases the fiber in a ferrule and holds optical fibers in face-to-face abutment with a threaded fiber locking mechanism is one possible connector. A high-precision ceramic ferrule connector with antirotation key may alternatively be used as the connector. This type of connector minimizes insertion losses when coupling single mode fibers. One such connector type shown in circle 7 is shown in the expanded cross-section in Fig. 7. With reference to Fig. 7, optical fiber 203 enters fiber connector output 214. In this example, fiber connector output is a ferrule. Fiber connector output 214 may be removably coupled to fiber connector input 215. Fiber connector input 215 extends through system casing 275. At the terminal end of the ferrule is fiber connector output 223. In a simple form, fiber connector output 223 may simply be the optical fiber end such that beam 237 is projected from the optical fiber.
Returning to Fig. 1A, fiber connector outputs 223, 222, 221 extend the optical fiber through the system casing 275 and project the light from their respective fibers onto respective collimating lenses 233, 232, 231 which collimate the output light into collimated beams 237, 236, 235 respectively. Beam 237 is directed by steering mirror 243 to dichroic mirror 242. Dichroic mirror is selected to allow the relatively longer wavelength light produced by laser 271 to be transmitted. While collimated beam 237 passes through dichroic mirror 242, collimated beam 236 generated by laser 272 is of a shorter wavelength that is reflected by dichroic mirror 242. In this way, collimated beams 237 and 236 may be combined. In a similar manner, dichroic mirror 241 transmits through longer wavelength light of collimated beams 236, 237 but reflects shorter wavelength light of collimated beam 235 thereby combining collimated beams 235, 236, 237 through use of a multi-mirror dichroic stack.
The collimated beams 235, 236, 237 are combined to form excitation beam 80. Excitation beam 80 impinges upon achromatic coupling lens 251 which focuses the beam into fiber connector input 253 on one end of optical fiber 255. Excitation beam 80 is transmitted through optical fiber 255 to fiber connector output 257 which projects the light from optical fiber 255 onto collimating lens 261. Optical fiber 255 is selected to have a fiber core sufficiently large or length sufficiently short to support the light of excitation beam 80 in guided mode. This light will be generally between 400 - 1000 nm wavelength light, typically from 600 - 700 nm wavelength if the system is being used to analyze whole blood. A small core, single, or multimode optical fiber accommodates this wavelength range.
Excitation beam 80 is collimated by collimating lens 261 and directed through glass plate 12 which optically communicates with power monitor 11. After passing through glass plate 12, excitation beam 80 then passes through laser line filter 13 and through spectral dispersion means 14 which acts as a mirror for the excitation beam wavelengths. Dispersion mirror 14 directs excitation beam 80 to mirror 15 which directs excitation beam 80 to the objective optics. The objective optics, power monitor, laser line filter and spectral dispersion means are all discussed more fully herein. The arrangement of Fig. 1A has the advantage that a user of the optical scanner system may connect and disconnect fibers coupled to different lasers such that any number of lasers may be used alone or changeably in different combinations. The system optics are pre- aligned and use of precision type fiber connectors allows for automatic fiber alignment.
The excitation beam 80 shown in Fig. 1, after being generated by the laser or lasers, passes first through glass plate 12. Glass plate 12 is in optical communication with power monitor 11. Those skilled in the art will realize several uses for power monitor 11 including ensuring proper laser output, detecting beam fluctuation or power loss, and to normalize later data by correcting for artifacts produced by beam power fluctuation.
After passing through glass plate 12, the excitation beam then passes through laser line filter 13 and through spectral dispersion means 14. A laser line filter is known in the art to reduce transmission of light that is not of the excitation light wavelength. Laser line filters include band pass interference filters and absorption filters. A laser line filter eliminates plasma tube light from the excitation beam 80, reducing optical noise in the system. This helps to ensure that light from the laser is not improperly directed by spectral dispersion means 14 to the detector. Spectral dispersion means 14 acts as a mirror for selected wavelengths. The spectral dispersion device may be, for example, a dichroic beam splitter, a prism, or a grating. The excitation beam is directed by spectral dispersion means 14 to mirror 15 and through right angle prism 16 to scan assembly 34. In Fig. 1, scan assembly 34 comprises a galvanometer 17 with attached galvo mirror 18, lenses 26 and 27, and lens 19. Alternatively, the scan assembly may comprise a multifaceted polygonal mirror. It is also possible to use a Bragg cell to produce the continual back and forth movement of the laser light. As was noted in the Background Art, U.S. Pat. No. 4,665,553, a Bragg cell may be used in optical scanners to produce the needed continuous back and forth scan of light. As illustrated in Fig. IB, a Bragg cell may be used to replace galvo mirror 18 and galvanometer 17 of Fig. 1. In this embodiment, fixed mirrors 61, 62, 63 direct excitation beam 80 to Bragg cell 65 at an angle θ relative to the optical axis. RF drive 67 introduces an acoustic sound field within the Bragg cell. This acoustic field causes the excitation beam 80 to be deflected through angle θ1. This deflection varies to produce a back and forth movement of the excitation beam along one axis. Returning to Fig. 1, the excitation beam 80 of the present invention impinges upon galvo mirror 18 which continually changes position because it is in communication with galvanometer 17 thereby causing a continuous change of position of the excitation beam across one axis. Thus the galvo mirror 18 in combination with the galvanometer 17 produces a continual back and forth movement of the excitation beam. Within scan assembly 34, the excitation beam travels from the galvo mirror 18 through lens 27 then through lens 26. From lens 26, the excitation beam is directed through lens 19 so that a focal spot of the beam may impinge upon the outer wall of a transparent container 20. The excitation beam impinging upon the outer wall traverses the wall and illuminates a columnar region of the sample causing fluorescent emission from the sample. Light collection may occur in an epi-illumination manner. It is also possible to invert the configuration and direct light into the container from the container's bottom through a transparent bottom well. The emitted fluorescence is collected by lens 19 and directed back, as retrobeam 83, through scan assembly 34. Lens 19, seen in Fig. 2, has a central portion for passage of incident beam 80 and uniform depth of focus of incident beam 80 into container 20. Because fluorescent emission is over a very wide angle, represented by rays 32a and 32b, fluorescent collection occurs over a wider portion of objective 19.
Returning to Fig. 1, the retrobeam 83 travels from scan assembly 34 to right angle prism 16 to mirror 15 and spectral dispersion device 14. Spectral dispersion device 14 is selected to reflect light of the wavelength of excitation beam 80 and transmit light of the wavelength of retrobeam 83. Thus, due to its fluorescence emission wavelength, retrobeam 83 is transmitted through spectral dispersion device 14 and through bandpass filter 21 to mirror 22 where it is directed through collimating lens 23. The retrobeam is then selectively passed through spatial filter 24 and into the detection means 35. The spatial filter 24 has a predetermined pinhole aperture of a diameter that permits passage of only that fluorescence emission from a region defined by the illuminated segment within the container as discussed below.
Detection means 35 may comprise a detection channel such as detector 30 which reads the fluorescent signal of the retrobeam 83 and is in communication with data reader 50 which converts it from analog to digital form. The detector is a light measuring device such as a photomultiplier tube or photodiode. The signal is recorded by data reader 50 as a unit of fluorescence intensity. The detection means 35 may contain any number of detection channels. For instance, a spectral dispersion device 25 is positioned between spatial filter 24 and detectors 30 and 31 in Fig. 1 to separate the wavelengths of the fluorescent emission of the sample and to selectively direct light below a determined wavelength to one detector and light above the determined wavelength to a second detector. As discussed previously, the spectral dispersion devices can be a prism, grating, or dichroic mirror. In this manner, multiple spectral dispersion devices and multiple detectors may be incorporated into the detection means for detection of fluorescence at different wavelengths from multiple fluorophores. In a similar manner, multiple lasers may be utilized for excitation of the sample at different wavelengths .
In Fig. 1, the spectral dispersion device 25 is shown as a dichroic mirror and the two detectors 30, 31 each detect one range of wavelengths. For example, wavelengths from below about 680 could be reflected by dispersion device 25 onto detector 31 while wavelengths above about 680 pass through dispersion device 25 into detector 30.
Figs. 1C and ID illustrate the use of a prism as the spectral dispersion device. A prism and a segmented detector become detection means 35. In Fig. 1C, retrobeam 83 passes through aperture 24 and onto refracting prism 71. Refracting prism 71 is composed of a selected material and is of a selected shape such that the index of refraction is higher for shorter wavelengths than for longer wavelengths. Refracting prism is selected and positioned specifically to disperse longer wavelength light (above 400 nm) into component wavelengths. The selection of different materials or shapes for prisms achieve varied optical effect. For example a constant deviation dispersion prism, such as the Pellin-Broca prism may be used. This prism has the advantage that the deviation does not depend on the index of refraction or wavelength. In a system using a constant deviation dispersion prism, particular wavelengths may be directed onto a detector by rotation of the prism. Refracting prism 71 divides retrobeam 83 into component wavelengths represented by 83a, 83b, 83c. These component wavelength light 83a, 83b, 83c impinge upon segmented photodetector 73. Segmented photodetector may be a multi-channel photomultiplier tube or charge- coupled device. This detector has a detecting surface 74 that detects the impinging light and produces an analog signal. For example in a multichannel charge-coupled device, detecting surface 74 is made of a self-scanning metal-oxide semiconductor. The signal from the segmented detector is transmitted to a data reader as in Fig. 1.
In Fig. ID a direct vision prism detector is shown. The use of a direct vision prism allows dispersion without deviating light at the central wavelength of the impinging light. Retrobeam 83 passes through aperture 24 and onto achromatic collimating lens 75. Collimating lens 75 collimates the light which then impinges upon direct vision prism 77. Direct vision prism is comprised of two or more prisms. Direct vision prism 77 as shown is comprised of a first prism 1 and a second prism 2. The prisms are selected to transmit light in the range of 400-1000 nm. First prism 1 is made of a higher dispersion material and second dispersion prism is made of a lower dispersion material.
The angles of surfaces a,b,c; the angle of the impinging light; and the material selected for prism composition will determine the amount of dispersion and the wavelength of the central undeviated light. In one embodiment 400-1000 nm light will be measured by the present optical system. The central undeviated light could be one of the selected fluorescent dye emission wavelengths such as 667 nm or 695 nm. The positions of surfaces a,b,c produce angles n and m which affect the wavelength of the central undeviated light as well as the dispersion angles of various wavelengths. This can be calculated with the formula: dD VN-
A2 = (N2-l) (V--V2)
A- = angle m A2 = angle n dD = dispersion
V- = reciprocal relative dispersion of optical material comprising first prism 1
V2 = reciprocal relative dispersion of optical material comprising second prism 2 N- = prism 1 index N2 = prism 2 index
Selection of prism composition material, prism angles and light angle allows selection of the wavelength of the central undeviated wavelength. This enables increased dispersion linearity and selection of wavelength range of detection.
Collimated retrobeam 83 passes through direct vision prism 77 and is refracted into representative light beams 83d, 83e, 83f. Central wavelength 83e is undeviated. The light then impinges upon detecting surface 74 of segmented detector 73. The signal from the segmented detector 73 is transmitted to a data reader as in Fig. 1.
One aspect of the present invention, the system depth of focus, is illustrated in Fig. 2. Spatial filter 24 is selected with a pinhole aperture that collects light over a large numerical aperture, but confines the depth of detection to a selected depth. The spot size of excitation laser beam 80 on the outside wall of capillary tube 20 is of a generally constant diameter, and may be chosen to provide uniform illumination along the depth dimension of the sample container such as a capillary tube.
It is possible to use a capillary as a sample container and have the entire 25-225 micron interior depth of the container receive uniform illumination from the waist of focused excitation beam 80. Alternatively, as Fig. 2 illustrates, the combination of objective lens, collimating lens and aperture allows for a selected depth of field. Thus it is possible to use a sample container that is deeper than the focal depth (i.e., depth in which the focused laser light produces uniform illumination) . Fig. 2A illustrates the use of such a container with the optical system of the present invention. Objective lens 19 focuses excitation beam 80 onto the surface of the container 9 such that a Gaussian waist of light impinges into the container. The waist of the beam illuminates a selected depth d. Outside of depth d the beam's energy rapidly falls off. At width w- the excitation beam 80 has a relatively high divergence angle such that the energy of the light wavefront at wx is distributed over a broad width with a gradual fall -off of energy. At width w2 the excitation beam 80 is much more narrowly focused with a greatly reduced divergence. The energy distribution along the beam width at the waist is much more uniform. Detecting within the beam waist allows uniform illumination of the volume from which fluorescence is measured. At width w3 the excitation beam again has a relatively high divergence angle and the energy of the light front is distributed over a broad, flattened bell curve. If a binding agent-binding site complex 45b is located in this site, the lower concentration of energy from the illumination light produces a greatly reduced amount of excited fluorescence from the complex.
In addition the combination of placement on lens 19 and the diameter of spatial filter 24 illustrated in Figs. 2, 2B will have an effect on the depth of field detected. This is governed by the formula
w0 = 4λf/πd where w0 is the spot size; λ is the light wavelength; f is the focal length of the lens; and d is the diameter of the input beam.
This formula for Gaussian laser beam propagation allows for calculation of beam spot size and design of system geometry. The formula provides an approximation of the spot size in a simple calculation.
As shown in Figs . 2 , 2A and 2B fluorescent emission is gathered over a wide angle. In relation to Fig. 2A representative fluorescent rays p,q,r emitted from binding agent -binding site complex 45a shown in Fig. 2A would impinge upon lens 19, and travel as collimated retrobeam 83 to the focus lens and detector. In contrast, binding agent-binding site complex 45b is outside of the focal depth of field. Although ray t may reach the detector, rays s,u will either not impinge on lens 19 or will be blocked by spatial filter 24. Thus in addition to weak illumination, the area outside of the depth of focus will have greatly diminished detectability.
It is known in the art to use an aperture as a field stop. This use of an aperture in combination with an imaging lens will restrict the effective aperture of the lens and reduce aberration. However a field stop could be ineffective to limit the depth of field, especially if the desire is to match the illumination volume to the collection volume. In contrast, the present invention uses a spatial filter to limit the depth of field. This spatial filter is positioned between the detector and the source of fluorescent emitted light. The spatial filter restricts the detection of light to rays of light emanating from a selected depth. This selected depth as shown is dependent on the geometry and placement of the spatial filter in relation to a lens. Matching the focal length of a lens to an aperture with specific geometry enables the present system to condition rays of emanating light to limit detection to a specific height of sample volume. The aperture is matched to the divergence of the lens to limit a depth of field. It is preferred that the spatial filter limit the depth of field to a depth of at least 25 nm. This use of a spatial filter is similar to the use of an aperture/focal lens combination in confocal microscopy to limit depth of field. The use of a confocal type aperture as the spatial filter of the present system enables "macro-confocal " scanning, i.e. the optical interrogation with a focused gaussian beam of excitation light with the optical interrogation limited by a spatial filter to a selected depth of field. The size of the aperture and the placement of the aperture in relation to the focus lens in combination with the properties of the focused gaussian beam will determine the selected depth of field. Thus the use of a spatial filter allows optical interrogation of a limited depth of detection in a variety of containers. In each container the depth of field is limited to a selected depth. This limited depth of field creates an optical interrogation depth that may be similar to the depth of a capillary. This "virtual capillary" detection allows the present optical system to optically interrogate a container wherein the container has a depth much larger than the detected depth of field. This limits the detected background fluorescent substantially to fluorescent light emanating from a specific depth of field. Fig. 2B illustrates how the size of the aperture can narrow the depth of field of the detection system. Spatial filter positioning indicated by arrows 124a results in a smaller aperture diameter such that additional light from the retrobeam is blocked. This means that the light from only retrobeam 183 will be detected by detector 31. This narrows the angle over which fluorescence is collected, represented by rays 132a and 132b. This effectively narrows the depth of field of the detector. In contrast, spatial filter positioning 24a results in a larger aperture diameter and less light from the retrobeam is blocked. Thus retrobeam 83 reaches the detector. This aperture geometry in combination with matching the aperture position to focal lens enables confocal-like narrowing of depth of field.
Fig. 2C illustrates the effect of objective lens selection on the shape of the beam waist . The shape of the beam waist determines the size of the beam spot and affects the optical depth of field. As was seen in Fig. 2, in Fig. 2C excitation beam 80 impinges upon lens 19 which focuses excitation beam 80 into a waist. Where the focused excitation beam impinges container 20 is indicated by spot width 31 (beam spot size) . The depth of focus is indicated by arrow dl .
Selection of different objective lens and different objective lens placement can alter the Gaussian waist and spot size. Lens 119 is shown positioned closer to container 20. Lens 119 is shaped such that narrow focusing of excitation beam 80 is achieved producing steeper divergence angles. A combination of lens shape and lens placement results in smaller spot size 131, narrower depth of focus indicated by arrow d2 , and a much greater divergence angle. The steeper divergence angle results in a more rapid falling-off of the beam's wavefield, i.e. the beam's energy. In addition the smaller size of lens 119 results in lesser retrieval of collimated light into a retrobeam. The capillary tube 20 of Fig. 2 may be a transparent sample holder of known dimensions . The capillary tube preferably has a rectangular cross-section with a shorter dimension defining an interior depth of 25 to 225 microns and a longer dimension defining a width of 1 millimeter. The beginning and ending points of the scan in a direction along the length of the capillary tube define the precise volume of the segment scanned. In the present invention, a capillary tube length of 40 millimeters has generally been used. In one embodiment, the capillary tube is fixedly positioned directly below the excitation beam so that the scan of the capillary tube occurs in a top-down manner. Alternatively, the optics may be reversed for bottom-up scanning. If a bottom-up scan is effected, the container scanned may be as deep or many times deeper than the depth of field or focal depth. If the container is a cuvette or microplate well, the liquid depth will generally be many times deeper than the depth of focus, with generally at least 25 μm used for focal depth. If containers such as microplate wells or cuvettes are used, various means may be used to focus the beam waist into the container. These methods are described in U.S. Pat. No. 5,556,764 and U.S. patent application Ser. No. 09/245,782, both hereby expressly incorporated by reference herein.
The intersection of the excitation beam and the container is generally defined by columnar region 51, the waist of the focused beam, as shown in Figs. 2 and 3. The top dimension of the columnar region is circular beam spot 33 which may be of 5 to 15 microns. Such beam spots produce illumination at least 25 nm deep. In one embodiment, the size of the beam spot is chosen so that the entire depth dimension of the capillary tube is illuminated by the waist of the focused beam.
As a columnar region of the capillary tube is illuminated, or the waist illuminates a focal depth within some alternate container, and the fluorescence emitted from its contents is detected and recorded, the optical scanning means (i.e. galvanometer, Bragg cell, rotating polygonal mirror) moves the columnar illumination to a new position to continuously illuminate new co- lumnar regions. The detection is affected by interval recording by the data reader. This ensures that only a fraction of the beam spot size is optically detected, measured and recorded at each measurement. The detection data capture is paced such that each illuminated columnar region detected partially overlaps another such region, as schematically illustrated in Fig. 3. The optical scanning and data collection continues in this manner of continuously illuminating and fluorescently exciting a columnar depth from which fluorescent emission is de- tected and periodically recorded and then repeating the process as the scan is effected.
In one embodiment, the optical scanning means follows a scan path in one direction indicated by arrow 52 that is transverse to the longitudinal axis of the capillary tube, i.e. along its width, and in the other direction along the length of the capillary tube, indicated by arrow 53. The paced data reads from the detectors to produce a two-dimensional array of beam spots as data. In Fig. 1, the dashed lines 134 indicate a change of position of scan assembly 34, so that dashed galvanometer 117 and dashed galvo mirror 118 represent galvanometer 17 and galvo mirror 18 in altered positions. In the same manner, dashed lenses 127, 126 and 119 represent lenses 27, 26, and 19 in altered positions. As shown in Figs . 3 and 4 , the transverse scan begins and ends at points 54 beyond the lateral boundaries of the capillary tube. This overscan is effective in identifying edge anomalies of the capillary tube.
Further description of data capture and processing is found in U.S. Pat. No. 5,556,764 hereby expressly incorporated by reference herein.
Fig. 5 shows a schematic representation of scan path 48 from above capillary tube 20, according to one embodiment of the present invention. The excitation beam spots are moved along the scan path in a transverse direction, then continuously snapped back to follow a closely-spaced parallel path also in the transverse di- rection. The process is repeated continually so that the scan also covers a segment along the longitudinal axis of the capillary tube. In this manner, fluorescence emission occurs and is detected from any chosen length of the capillary tube. The method disclosed in the present invention allows for analysis of a sample of biological fluid with a minimum of preparation. According to the present invention, a fluid is incubated with an excess amount of a binding agent that contains a fluorophore of known optical characteristics. The fluorescently-labeled binding agent is selected to react with binding sites present within the sample. For example, a fluorescently-labeled antibody directed to an antigen present on some cellular component of a biological fluid may be added to the sample. The labeled binding agents and the binding sites form fluorescent complexes that will emit a signal when analyzed by the apparatus of the present invention. After the sample is incubated with the labeled binding agent, it is diluted, if necessary, and then placed directly into capillary tube 20 or other container. No lysing of components of the biological fluid nor separation of bound and unbound binding agent is required at any point in the practice of the method of the present invention. An optical scan is made of the sample in a volumetric manner and fluorescence emission is continuously sequentially recorded from overlapping columnar region.
The enumeration may occur in an absolute volume, depending on the desired application, by noting the beginning and ending points of the lengthwise scan of the capillary tube and measuring the distance scanned or by scanning between specific identification marks on the container. This quantitation of all of the fluorescent complexes in a fixed, precise volume is a powerful method of quickly obtaining detailed population data. This volume may be fixed by using a uniform cross-sectional area capillary tube or by independently measuring the volume of the capillary tube between specific identification marks or by simply noting the length and width of a scan of a known depth of field. Alternatively, a ratio can be obtained without counting a precise volume, but rather by comparing relative counts of different components of the sample.
Data reader 50, in Fig. 1, records events, i.e. an increase above the background level of fluorescence exceeding some threshold value, as shown in Fig. 6. The events correspond to the occurrences of cells of a particular type in the sample. Fluorescence emission occurs from both the binding agent-binding site complexes 45 and from the free binding agent 40, but a more intense signal 85 relative to background level 80 comes from areas where the binding agent is clustered, i.e. cells or other targets exhibiting binding sites to which the binding agent is directed. Therefore, a signal of heightened fluorescence 85 corresponds to a cell, linked by dashed lines in Fig. 6, and is recorded as such by an electronic data processor of the data reader. The data acquisition and analysis may be effected in the manner described in U.S. Pat. No. 5,556,764 hereby expressly incorporated by reference herein.
The method of the present invention does not require removal of unreacted fluorescently-labeled binding agent. Dilution of the sample before optical scanning serves to improve signal to noise ratios so that fluorescent imaging according to the present invention occurs in a rapid manner with minimal processing steps. Dilution also serves to minimize the occurrence of cell overlap in the sample. In the practice of this invention, optimal results are obtained when the cells of the sample are on the order of 10 microns in size and the cell density is less than 5000 per microliter.
An example of this method's utility is illustrated by an assay for the determination of leukocyte subclasses present in a blood sample. In a typical assay, a sample of whole uncoagulated blood is incubated with fluorescently-labeled antibodies that are directed to specific cell surface markers. For example, anti-CD4 and anti-CD8 labeled with fluorophores having different optical characteristics may be incubated with the whole blood sample. Within the sample, the leukocytes that bear either CD4 or CD8 or both cell surface markers will react with the labeled antibodies to form fluorescent complexes. After a sufficient reaction time, the whole blood sample is diluted and inserted into a capillary tube. The capillary tube is then optically scanned according to the present invention. The wavelength range of the optical scan is selected so as to activate the fluorophores being used and to minimize interference due to autofluorescence from blood components . Fluorescent emission corresponding to the fluorophores that were used to label the anti-CD4 and anti-CD8 are detected and recorded. The presence of leukocytes bearing either or both of the cell surface markers to which these antibodies are directed is then enumerated. Results may be presented as an absolute cell count per unit volume, by counting the number of cells of a certain subclass present within a given volume, the volume being determined by the length of the scan and the cross-sectional area of the capillary tube. Results may also be presented as ratios, e.g. CD4/CD8 leukocyte ratios, by counting the number of cells bearing each of these cell surface markers and comparing the two. The usefulness of this last illustration is readily evident, as this ratio is important in determining the progression of AIDS. This assay is further described in U.S. Pat. No. 5,585,246, hereby expressly incorporated by reference herein. When an assay is performed to determine leukocyte subclasses in whole uncoagulated blood using the technique of the present invention, a two or three minute wait between placement of the reacted sample into the capillary tube and the optical scan allows for the natu- ral density of the numerous red blood cells present in the sample to cause settling of the red blood cells to the bottom of the capillary tube and the subsequent displacement of the white blood cells. This natural buoyan- cy effect causes a resultant location of the white blood cells near the upper portion of the capillary tube and assists in fluorescence detection in a top-down scan geometry. As in the above example, fluorophores with different optical characteristics can be combined with binding agents directed to different binding sites, so that the presence of multiple targets in the sample can be detected. From the precise known volume of the container that has been scanned, a quick reading will identify the number of cells of a particular subclass per unit volume that are present in the sample. The optical system is simply set to excite each fluorophore at its excitation wavelength and a detection channel or other noted detection optic is created to correspond to the emission wavelength of each fluorophore.
The apparatus and method of the present invention are suited to many applications, including those requiring absolute counts of complexes within a known volume. For example, cell kinetics studies, cell toxicity studies using intercalating dyes, receptor binding assays and in situ hybridization may be adapted for analysis according to the present invention. The apparatus is directed to detect fluorescence in a limited depth of field which could be used for detection of fluorescent reporter beads, observing fluorescent compounds produced by cells or any other application where fluorescence may be observed in a limited depth of field. Because the level of fluorescence is measured, the amount of binding sites may be determined. In addition, the volumetric method of the present invention allows for the avoidance of artifacts that may be present when immunological and other biochemical responses are studied in cells located on a surface rather than in a cell suspension. Although the cells are detected from within a static sample container, the present invention presents the sample in a manner that allows for flow cytometric-type analysis on relatively stationary localized cells. Therefore, the cells may be detected in a location-specific manner or be identified for subsequent visual examination.
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|International Classification||G01N15/00, G01N21/64|
|Cooperative Classification||G01N21/645, G01N21/6428, G01N2021/6439, G01N2015/0065, G01N21/6458, G01N2021/6463, G01N21/6456|
|European Classification||G01N21/64H, G01N21/64P|
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