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resolution color CCD cameras are typically below 2 Hz, which is too slow for scanning around on a sample and makes focusing difficult. With these cameras it is necessary to focus visually through the microscope, then switch the light path to the camera and refocus and readjust the illumination. The process requires operating two instruments to get a digital image. The large image sizes of some cameras (viz. 3200 2400 pixels) creates an additional problem. The amount of data in one image can be so large that the time to transfer those data from the camera to a computer for display is significant, taking up to 1 min in some cases. Strategies to increase the frame rate for scanning and focusing include transmitting only a small fraction of the total image. The resulting tiny, low-resolution black-and-white images are not ideal for scanning or focusing. Therefore, the time required to establish the best focus, color levels, and intensity levels and store the image can be very long indeed.

In this scanning microscope the image has been standardized at 1280 1024 pixels, which means that the resolution is limited by the Airy disk size of the objective and not by the electronics or pixel size of a detector. With the scan rate of 13.4 Hz (frames per second), it is easy to focus and move the image. Indeed, one can obtain a large number of good images in a short time using this system—a major advantage from a user perspective. There is no reason why a larger image size could not be used, with the caveat that higher file size and slower image frequency would result. Of course, these factors are frequently based upon current computational capabilities, a factor that changes constantly.

There are several other significant differences between COSMIC and traditional camera-based systems. Most, if not all, of the following issues are not applicable to the white-light scanning microscope but may be significant problems for camera systems under certain operating conditions.

First, the regular array pattern of pixels in a CCD camera can cause the generation of false data. When the sample has a periodic structure whose spacing is about the same as that of the detector pixel spacing, a beat frequency can be generated from slight misalignments between the image and the detectors. The result is wavelike patterns appearing in the picture. These patterns, called aliases, or Moiré patterns, are not real in the image but are caused by the interaction of the image with the detector. These aliasing patterns can sometimes be eliminated by rotating the image against the detector. Aliasing can be a significant problem for certain kinds of images, particularly for automated image analysis. Aliasing problems can be greater for color cameras that have three CCD detectors, one each for the red, green, and blue portions of the light spectrum. The three detectors must be carefully aligned to give the same size image for each color and to prevent aliasing between the detectors.

Second, if a portion of a CCD detector is saturated with light, the excess electrons or the excess light may cause a signal from adjacent pixels. The visual effect may look like a star pattern with perpendicular arms and is called blooming. Electronics to detect and minimize blooming are usually built into the camera; however, the result is that the operation of the camera is temporarily halted.

Third, in an ordinary microscope, all points in the image field are illuminated simultaneously, which is called full-field illumination. In this scanning microscope, only a single point is illuminated at a time since illumination is provided by a flying

Operational Considerations

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spot. When all the pixels are simultaneously illuminated, scattered light can reduce contrast from pixel to pixel and contribute to blooming.

The fourth issue is interpolation. Medical systems require better resolution than can be obtained from inexpensive mass-produced television-format CCD cameras. Some camera manufacturers have chosen to create artificial resolution by calculation. Several versions have been implemented. One method is to increase resolution simply by interpolation between pixels. In another case, one of the three color detectors is displaced one-half pixel horizontally and one-half pixel vertically. The color data for that physical location are computed from the nearest neighbors. In this technique, there are no physical locations with true information in any part of the image since all data must be computed. It is not clear how interpolation impacts the use of these detector systems, but the issue will have to be addressed eventually. Finally, because the sensitivity of PMTs is at least 1000 times greater than typical CCD detectors, the illumination intensity of the white-light scanning microscope can be 1000 times less than traditional light sources, reducing bleaching or photodamage to the specimen, an issue frequently overlooked in biological imaging.

OPERATIONAL CONSIDERATIONS

The most significant problems to be dealt with in a CRT-based spot-scanning instrument are the cost of the instrument and the aging characteristics of the CRT, which will decrease in light output over time. CRT lifetimes in the 10,000-h range are common, and replacement is not difficult or expensive, particularly when compared with most lasers used in microscopy. However, extended use of the zoom feature can cause preferential aging of the center of the CRT screen. Therefore, zoom time is limited by a timer circuit that disables the zoom after a preset interval (2 to 10 min.), thereby preventing the zoom feature from aging the center of the screen. The brightness and color balance calibration procedure continually corrects for CRT aging until the overall light level is unacceptable and a replacement is needed. Changing or adding different objectives requires calibration for that objective, but once performed, the calibration becomes an instrument preset.

A troublesome issue for those who must deal with microscope and camera systems is the difficulty of obtaining rapid and reproducible images from camera/computer systems. Combining the camera, imaging boards, and computer software is, in itself, an interesting challenge. In this system there are fewer variables and therefore a more consistent and stable image platform.

The clearest departure from traditional microscopy in this scanning microscope is the lack of a viewing ocular. This is a radical difference for microscopists who may feel uncomfortable not being able to view the specimen, except electronically. However, the format is most likely consistent with the future of digital imaging applications. Clearly, the whole concept of telepathology or remote evaluation of images implies that at least one participant will not have access to a traditional viewing ocular. Thus, the complete removal of the ocular may actually be advantageous for the originator of the images to know exactly what the remote viewer is seeing.

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The COSMIC instrument is clearly the first fully integrated digital microscope system to offer real-time, full-color, high-resolution digital images. The unique features of zoom magnification, contrast enhancement and suppression, averaging and summing, background correction, and superresolution capabilities are major advantages. The lack of moving parts and alignment protocols is a powerful incentive for use in service roles such as pathology, where a robust, easy-to-operate, high-resolution instrument is desired.

FUTURE APPLICATIONS

Because of the nature of the optical pathway, transmitted fluorescence detection is likely to be a viable option using this technology. Currently, no commercial microscopes use transmitted fluorescence; they are all epifluorescent systems. Because bright-field and fluorescence in a simple, easy-to-use system is a desirable combination, addition of fluorescence to this scanning microscope will significantly enhance its potential in pathology.

REFERENCES

Bartels PH, Layton J, Shoemaker RL (1984): Digital microscopy. Monogr Clin Cytol 9:28–61.

Bibbo M, Bartels PH, Dytch HE, Puls JH, Pishotta FT, Wied GL (1983): High-resolution color video cytophotometry. Cell Biophys 5:61–69.

Ingram M, Preston K, Jr (1970): Automatic analysis of blood cells. Sci Am 223:72–82.

Ingram M, Preston K, Jr (1964): Importance of automatic pattern recognition techniques in the early detection of altered blood cell production. Ann NY Acad Sci 113:1066–1066.

Mellors RC, Keane JF Jr, Papanicolaou GN (1952): Nucleic acid content of the squamous cancer cell. Science 116:265–265.

Nagata H, Mizushima H (1998): World wide microscope: new concept of internet telepathology microscope and implementation of the prototype. Medinfo 9 Pt 1:286–289.

Plasek J, Reischig J (1998): Transmitted-light microscopy for biology: A physicist’s point of view. Proc RMS 33:196–205.

Ploem JS (1988): Modern image analysis methods in hematology. Nouv Rev Fr Hematol 30:45–49.

Ploem JS, van Driel-Kulker AMJ, Goyarts-Veldstra L, Ploem-Zaaijer JJ, Verwoerd NP, van der Zwan M (1986): Image analysis combined with quantitative cytochemistry. Results and instrumental developments for cancer diagnosis. Histochemistry 84:549–555.

Prewitt JM, Mendelsohn ML (1996): The analysis of cell images. Ann NY Acad Sci 128:1035–1053.

Singson RP, Natarajan S, Greenson JK, Marchevsky AM (1999): Virtual microscopy and the Internet as telepathology consultation tools. A study of gastrointestinal biopsy specimens. Am J Clin Pathol 111:792–795.

Szymas J, Wolf G (1998): Telepathology by the internet. Adv Clin Path 2:133–135.

Vazir MH, Loane MA, Wootton R (1998): A pilot study of low-cost dynamic telepathology using the public telephone network. J Telemed Telecare 4:168–171.

Emerging Tools for Single-Cell Analysis: Advances in Optical Measurement Technologies

Edited by Gary Durack, J. Paul Robinson Copyright © 2000 Wiley-Liss, Inc.

ISBNs: 0-471-31575-3 (Hardback); 0-471-22484-7 (Electronic)

14

Illumination Sources

Howard M. Shapiro

Howard M. Shapiro, M.D., P.C., West Newton, Massachusetts

REQUIREMENTS FOR A LIGHT SOURCE FOR CYTOMETRY

Cytometric measurements involve the collection of light from small regions of space for short periods of time. In a flow cytometer, measurements of light scattering, transmission, and fluorescence are made in a few microseconds; while longer, the measurement interval in imaging or scanning static cytometers is still short compared to the time needed to examine cells under a microscope. Since even the most efficient light collection optics capture photons over a relatively small solid angle, and there are typically substantial losses associated with wavelength selection and detection, light sources used for cytometry must emit a relatively large number of photons in a desired wavelength region per unit time, and it must be possible to direct a substantial fraction of this photon flux into a small volume of the sample. Two classes of light sources best meet these criteria; they are so-called short or compact arc lamps and lasers.

RADIOMETRIC UNITS AND LIGHT SOURCE CHARACTERISTICS: EXITANCE, RADIANCE, AND THROUGHPUT

The process by which radiant energy or radiation is measured, with all wavelengths given equal weight, is defined as radiometry; in the International System of Units (SI

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Units), radiant energy is expressed in joules and radiant flux (energy per unit time) in watts, where one watt equals one joule per second.

Radiant areance, or exitance, expresses the power or radiant flux ( ) emitted per unit area (A), in units of watts per square meter. A true point source could be described in terms of its intensity, or pointance, that is, the power emitted per unit solid angle (ω), in units of watts per steradian. (A sphere of radius r has a surface area of 4 r2; one steradian is defined as that solid angle intercepting an area equal to r2 on the surface of the sphere.) Most of the light sources encountered in the real world are extended sources such as arc and filament lamps and the sun and other stars, which, while they may be capable of emitting substantial amounts of power, disperse their emission in all directions, that is, over a solid angle of 4 steradians. Their behavior is characterized in terms of radiance, which is the power emitted from, transmitted through, or reflected by a surface per unit of its area per unit solid angle. The units of radiance are watts per square meter per steradian.

The optical system that must be placed between an extended source and a specimen must, at a minimum, include a lens to collect light from the source and another to converge the collected light on the specimen. Both lenses are generally referred to as condensers; for purposes of this discussion, they will be distinguished as the lamp condenser and the microscope condenser. To add to the confusion, it should be noted that, in the epi-illumination systems generally used in fluorescence microscopes, the microscope objective serves as the microscope condenser.

The most efficient illumination, that is, that which produces maximum photon flux through the specimen, can be obtained by making an image of some portion of the emitting surface in the plane of the specimen. In order to maximize light collection and specimen illumination, the condenser lenses should be of high numerical aperture (NA) (equivalent to a low f number) to make the collection and illumination angles as large as possible. Consider, for the moment, that a single convex lens serves as both lamp condenser and microscope condenser. If the emitting surface of the lamp is located two focal lengths away from the lens on one side, and the specimen plane is located two focal lengths away on the other, the lens will form a 1 : 1 image of the source in the specimen plane, and the solid angles for collection and illumination will be the same. If a magnified image of the source is to be formed, the lamp must be placed closer to the lens, and the specimen moved further away; the lens can then collect light over a larger solid angle, but the photon flux per unit area in the specimen plane will not be increased. If a minified image of the source is formed, the lamp must be placed more than two focal lengths away from the source and will therefore collect light over a smaller solid angle; again, the photon flux through the specimen will not be increased. A quantity called optical throughput is said to be conserved; no matter how a real optical system is arranged, there is a point beyond which photons can be lost but not gained.

The largest possible collection angle obtainable using a convex collecting lens will be achieved if the lens is placed one focal length from the source; it will then collect a collimated beam of light, that is, one in which all the “rays” are parallel, requiring that a second lens be used to converge light on the specimen. It is also possible to col-

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lect light from a substantial fraction of the emitting surface of a source using ellipsoidal or parabolic reflectors. However, while this can allow substantial amounts of power to be directed through surface areas larger than the emitting surface of the source, it is still not possible to put more photon flux per unit area of the specimen than can be collected per unit area from the source.

LAMP AND LASER CHARACTERISTICS COMPARED

Figure 14.1, which originally appeared in a catalog from Oriel Corporation (Stratford, CT; www.oriel.com), a major supplier of light sources, compares the irradiance of mercury (Hg) and xenon (Xe) arc and quartz tungsten halogen (QTH) filament lamps. The irradiance values are given in milliwatts per square meter area per nanometer wavelength at a distance of 0.5 m; the corresponding radiance is calculated by determining the solid angle represented by a square meter at this distance and by integrating over an appropriate wavelength region.

The 100-W Hg arc lamp has an irradiance of about 120 mW m-2 nm-1 in a 10-nm band around the strong ultraviolet (UV) line at 366 nm, so total irradiance in this spectral band is approximately 1.2 W m-2. The surface area (4 r2) of a sphere with a radius of 0.5 m is 4 (0.25) m2. Since this is the surface area subtended by a solid angle of 4 steradians, 0.25 m2 is the surface area occupied by 1 sr; the radiance is thus about 0.3 W m-2 sr-1, or 300 mW m-2 sr-1.

The arc in a 100-W Hg lamp is 0.25 mm (250 µm) in diameter; its surface area is 4 (0.000125)2 m2; therefore the surface area of a 1-sr segment of the arc is (0.000125)2 m2, or 15625 µm2. The power radiated through 1 sr is 300 mW. A spot 10 µm in diameter on the surface has an area of (5)2 µm2, or 78.5 µm2; thus, about 1.5 mW [(78.5/15625) 300] of UV power in the bandwidth discussed would be emitted through this surface. This represents the maximum amount of power collected from the arc that can be directed back through the same area; thus, if the arc is used to illuminate a cell 10 µm in diameter, no more than 1.5 mW can impinge on the cell at any given time, no matter how efficiently light is collected from the arc and transmitted through the optics.

Mercury arc lamps have strong emission lines at several wavelengths other than 366 nm; peak radiance near that of the 366-nm line is also obtainable at 313 nm, farther in the UV, at 405 nm (violet), 436 nm (blue-violet), 546 nm (green), and 578 nm (yellow). Under the best conditions, neglecting transmission losses in lenses and in the filters used for wavelength selection, one would expect to be able to direct at most 1–1.5 mW in any of these wavelength regions from an Hg arc lamp through a 10-µm cell at any time.

A xenon arc lamp, as can be seen from Figure 14.1, has a relatively flat emission spectrum between the near ultraviolet and the near infrared (350–750 nm); this makes the Xe arc a desirable illumination source for spectrophotometers and spectrofluorometers. However, since the radiance of the Xe arc over this range is only about 1/10 the radiance of the Hg arc at its strong emission lines, the Xe arc is less desirable as

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a source for cytometry, except possibly in the region between 450 and 500 nm, where the radiance is slightly higher than that of Hg lamps.

The radiance of a QTH filament lamp is substantially lower than that of a Xe arc lamp below 600 nm and slightly higher between 600 and 800 nm. However, the area of the emitting surface of the filament lamp is much larger than the area of the arc in an arc lamp; thus, much less power can be collected from and directed through a small area. This makes filament lamps poorly suited for fluorescence excitation in flow cytometry, although they have been used quite successfully in flow cytometers that measure absorption and light scattering. For these purposes, they have the advantage that it is relatively easy to achieve precise regulation of output power using a well-regulated dc supply.

Although higher power Hg and Xe arc lamps emit more radiant energy than does a 100-W Hg arc lamp, they are less desirable as sources for cytometry because their arcs have larger surface areas, with the result that their emission per unit area is lower; such lamps may, however, offer an advantage in apparatus in which an area substantially larger than a millimeter or so in diameter must be illuminated, for example, a static cytometer with a charge-coupled device (CCD) or other imaging detector and a large field of view.

The effective power available from a 100-W Hg arc has been increased experimentally by Steen and Sørensen (1993), who modified the front-end electronics and arc-lamp power supply of a flow cytometer to increase the lamp current by a factor of 10 for a few microseconds after the trigger signal indicates the presence of a cell in the observation region. The modification is inexpensive, increases measurement sensitivity, and does not substantially decrease the life of the lamp but has not, to date, been incorporated into a commercial apparatus.

F i g . 14.1. Irradiance of mercury and xenon arc and TOH filament lamps. (From Shapiro HM (1994): Practical Flow Cytometry, 3rd ed. New York: Wiley-Liss. Courtesy of Oriel Corporation.)

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Laser sources differ radically from extended sources in several respects; one is that the emission from lasers is confined to a very small solid angle; it is generally possible (neglecting transmission losses) to focus all the energy in the beam to a circular or elliptical spot. Most lasers used for flow cytometry emit a beam in which the energy distribution is normal, or Gaussian. Spot “diameters,” in the case of a circular spot, or “width” and “height,” in the case of an elliptical spot, define the “1/e2 points,” the region within which approximately 87.5%, or 1 – (1/e2), of the total emission is contained. The area of this region corresponds to the area of a central elliptical region of a bivariate normal distribution within two standard deviations of the bivariate mean.

If a laser beam is focused to an elliptical spot 100 µm wide by 20 µm high, a not unreasonable size to use in a flow cytometer, it can be calculated that about 8.8% of the total power in the beam would illuminate a cell 10 µm in diameter at or near the center of the focal spot. If the 1.5 mW already shown to be the maximum available from a 100-W Hg arc lamp were to come instead from a laser, the total laser power would therefore be 1.5/.088 mW, or 17 mW. Assuming that wavelength selection filters, which must be used with an arc lamp but are generally not necessary when laser illumination is used, reduce light transmission by 25–30%, the arc lamp is, at best, equivalent to a laser emitting about 12 mW.

To summarize: If lamp illumination is to be used, the highest intensities, comparable to the output from a laser emitting no more than 12 mW, can be obtained from a 100-W Hg arc lamp (arc diameter approximately 0.25 mm), in spectral regions around 313, 366, 405, 436, 546, and 578 nm. In the region between 450 and 500 nm, slightly more power may be available from a 150-W Xe arc lamp than from a 100-W Hg arc lamp, but this will be equivalent to no more than 1–1.5 mW from a laser (e.g., an argon-ion laser emitting at 488 nm). Even less power is available from filament lamps; while these are usable and may even be desirable for measurements of absorption and light scattering because it is easy to regulate their output, they are poorly suited for fluorescence excitation. Lamps are generally preferable to lasers when uniform illumination over a relatively large area is needed, as is the case in many imaging cytometers.

NOISE IN LIGHT SOURCES AND ITS EFFECTS ON MEASUREMENTS

The output of a light source used for cytometry should, ideally, be constant; since, in the real world, there will be shortand long-term variations in power output, it is desirable to know the magnitudes of such variations, which will allow determination of their effects on measurements. In some cases, it is possible to achieve acceptable measurement precision using an otherwise unacceptably noisy light source by measuring temporal and spatial fluctuations in output and compensating for them using analog and/or digital electronics.

Both lamps and lasers typically exhibit drift in output power over periods of hours to days and power loss over a period of weeks to years due to aging and wear of various components of the source. Such long-term variation generally does not affect

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measurements made within a few hours of one another; if it is pronounced, it is generally dealt with either by interspersing standard samples at intervals during a day’s run or by repair or replacement of the offending source.

Shorter-term fluctuations, ranging in frequency from a few hertz to a few megahertz, can interfere with both measurement precision and measurement sensitivity. The noise of a light source is usually specified in terms of a percentage that can represent either a peak-to-peak or an RMS (root-mean-square) value. The RMS noise value is equivalent to the coefficient of variation (CV) of the fluctuations of power output about the mean. The RMS noise can be estimated from the peak-to-peak noise level by assuming that the peak-to-peak value represents a range of six standard deviations about the mean, the rationale being that more than 99% of the area of a normal distribution lies within three standard deviations of the mean. Thus, dividing the percentage of peak-to-peak noise by 6 provides an approximation to the RMS noise percentage, which, since it represents the CV of power fluctuations, is 100 times the standard deviation of the output power divided by the mean output power.

In flow cytometers, the dc component and slowly varying components of the noise are typically removed by baseline restoration circuits in the preamplifier electronics; the higher frequency components of the noise are superimposed on the output signal. When light scattering is measured, most of the background comes from stray illuminating light; since this is at the same wavelength as the light scattered from the cell, it cannot be reduced or removed by optical filters. When fluorescence is measured, the source-dependent component of the background typically comes from fluorescence of dye in the sample stream and from Raman scattering by water in the sheath and sample streams; potential contributions from illumination light leakage and filter fluorescence can and should be avoided by proper choices of dichroics and detector filters. The amplitude of the fluctuations contributes to the variances of both light scattering and fluorescence measurements, with proportionally greater effects on smaller signals; when the amplitudes of noise fluctuations and signals from particles become comparable, signals from the particles become undetectable, and signals from particles may become unreliable for triggering well before this point is reached.

Under the best of circumstances, the CV of a light scattering or fluorescence measurement cannot be lower than the RMS noise of the light source, except that the effects of source noise on fluorescence signals may be mitigated if the illumination power level is high enough to produce saturation in the relevant fluorophores (van den Engh and Farmer, 1992). For example, chromosome sorters that use relatively high (hundreds of milliwatts) UV and blue-violet high excitation power from ion lasers often achieve fluorescence measurement CVs on the order of 1% when RMS source noise is somewhat higher.

Because cells, even when stained, do not absorb more than a small fraction of the light passing through them, illumination fluctuations have larger effects on absorption and extinction measurements than on measurements of scattered light and fluorescence. If an extinction measurement is made of a particle that removes 10% of the light from the incident beam, a 1% increase in illumination intensity occurring while the particle passes through the beam results in an apparent light loss of 9%, while a

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1% decrease in illumination intensity results in an apparent light loss of 11%. Thus, a 1% change in illumination intensity produces a 10% change in the amplitude of an absorption or extinction signal; the less light the particle removes from the beam, the larger the effect of source intensity fluctuations. While the 1% RMS noise level of the air-cooled argon and He–Ne laser sources most commonly used in fluorescence flow cytometers does not significantly compromise the precision of scattering and fluorescence measurements, extinction measurements made using these sources are unacceptably imprecise.

In many cases, the noise level of both lamp and laser sources is determined primarily by the degree of regulation built into the power supply by the manufacturer and is therefore not controllable by the user. Although it is possible, at some sacrifice in power, to decrease output fluctuations using devices incorporating acousto-optic or electro-optic modulators and appropriate feedback circuitry, these typically cost several thousand dollars and require some technical proficiency to adjust. It is less expensive to add noise compensation circuits that monitor source output and adjust amplifier gain using analog difference or ratio circuits; these are not typically implemented in commercial flow cytometers. In the near future, it can be expected that both baseline restoration and source noise compensation will be facilitated by the use of digital signal processing, allowing pulse characteristics to be determined rapidly and precisely after signals from the detectors are digitized at frequencies of several megahertz.

Both lowand high-power ion lasers can generally be operated in either a current control mode or a light control mode. In the current control mode, the laser power supply is regulated to deliver a constant current, and light output remains constant only if the mechanical and optical characteristics of the laser do not change during operation. If if the laser becomes slightly misaligned, light output decreases, even though power supply current remains the same. In the light control mode, power supply output is regulated by a feedback circuit that samples the energy in the beam and adjusts the laser current to maintain constant light output. This generally works well when the laser is emitting at a single wavelength; however, when several lines are emitted simultaneously, it is more difficult to keep power constant. The air-cooled argon lasers in benchtop flow cytometers are operated in the light control mode; so are diode lasers, which are built with a light-sensing photodiode and need a light control feedback loop in the power supply to prevent thermal runaway from destroying the laser itself when it is turned on. Helium–neon and helium–cadmium lasers typically do not incorporate light control circuits. Operation in the light control mode does not guarantee protection against noise if the light output drops substantially enough for the supply current to rise to its maximum value; under these conditions, any further mechanical or optical deterioration cannot be compensated for by either light control or current control circuitry.

Arc lamps exhibit a phenomenon known as arc wander, in which the position of the arc between the electrodes moves over time; this results in changes of illumination intensity in the field of observation. In the author’s experience, arc wander has been associated with inadequate cooling of the lamp and housing more than with any other characteristic of the lamp or power supply. To compensate for arc wander, it is