Photodiagnosis and Photodynamic Therapy (2007) 4, 184—189
available at www.sciencedirect.com
journal homepage: www.elsevier.com/locate/pdpdt
Tetherless fiber-coupled optical sources for extended metronomic photodynamic therapy Nick Davies BSc (Hons), MSc a,∗, Brian C. Wilson a,b a
Department of Medical Biophysics, University of Toronto, 610 University Avenue, Toronto M5G 2M9, Canada Ontario Cancer Institute, University Health Network, 610 University Avenue, Toronto M5G 2M9, Canada Available online 8 June 2007
b
KEYWORDS Metronomic photodynamic therapy; Brain tumors; Optical sources; Light-emitting diodes; Multiport light delivery
Summary Background: Metronomic photodynamic therapy (mPDT) involves the delivery of a low fluence rate of light and photosensitizing drug, continuously and over an extended period. However, in mPDT trials, there has been a perceived need for light sources that can deliver light reliably for extended periods, while being small enough to be tolerated by small animals. Methods: We report on the development of tetherless, lightweight, fiber-coupled optical sources for in vivo delivery of interstitial mPDT. Two forms are reported, based on diode lasers and light-emitting diodes (LEDs), the latter types weighing only 16.5 g. Results: The prototypes have been well tolerated in preliminary trials on tumor-bearing rat models and can currently provide stable levels of performance for upwards of 5 days. We also report an extension of the concept to units coupled to several optical fibers to enable simultaneous irradiation of multiple locations in tissue. Conclusion: These prototypes fulfill a current need for reliable mPDT optical sources for use with animal models. They also serve as the foundation for the development of fiber-coupled sources for use in future clinical trials. © 2007 Elsevier B.V. All rights reserved.
Introduction Photodynamic therapy (PDT) is a minimally invasive form of cancer treatment that is dependent on light, a photosensitizer (PS) and oxygen in order to effect tumor cell death [1]. Conventionally, PDT is given in the form of a single, acute treatment (aPDT). In this form, the PS is delivered as a single bolus. Several hours post-drug administration, the tumor region is briefly irradiated using high-intensity light, resulting in tumor necrosis. However, aPDT treatments may also
∗
Corresponding author. Fax: +1 408 453 9293. E-mail address: [email protected] (N. Davies).
cause sensitization of normal tissue that is adjacent to the tumor and also exposed to the light field [2]. In particular, for PDT of malignant brain tumors, we have shown that damage to normal tissue, especially white matter, can be avoided and tumor cell-specific apoptotic cytotoxicity can be achieved by using aminolevulinic acid (ALA)-generated protoporphyrin IX (PpIX) if the ALA and light are delivered at low rates over an extended period [3]. This approach has been termed ‘metronomic PDT’ (mPDT) by analogy with metronomic chemotherapy, and is currently being optimized in pre-clinical rodent models bearing intracranially implanted malignant brain tumors, prior to planned clinical trials. Delivery of ALA in these animal models over an extended period is feasible by putting ALA into the drinking water, and we have shown that adequate
1572-1000/$ — see front matter © 2007 Elsevier B.V. All rights reserved. doi:10.1016/j.pdpdt.2007.03.005
Tetherless fiber-coupled optical sources for extended metronomic photodynamic therapy levels of PpIX can be achieved without significant toxicity [4]. Corresponding metronomic light delivery is more challenging, in that it is necessary to provide a means to generate and deliver the light to the intracranial tumor in a sterile, minimally invasive manner over at least several days and to administer an adequate total light fluence while also allowing the animals free movement. We have previously reported the development of both superficial and interstitial sources for mPDT delivery and the limitations associated with each [4]. In one model, the implanted optical fiber was coupled to a fixed, externally powered diode laser and the intervening length of fiber was mounted in an overhead ‘trolley car’ jig. To overcome the resulting limitations on animal mobility, we subsequently developed a battery-operated, tetherless light-emitting diode (LED) ‘backpack’. This weighed only 23 g and allowed the animal freedom of movement throughout the mPDT procedure. A long-life battery and potentiometer were housed in 5 cm × 3.5 cm × 2 cm box. The LED was placed directly into the cranium and fixed in place with bone cement over a 1 mm hole drilled in the skull. Powered by a 1.6 A h Keeper II Magnum battery (Allied Electronics, Fort Worth, TX), this device could deliver an output profile ranging from an initial power of 1.6 mW to a terminal value of 116 W over a period of 10 days. Although it was reasonably well tolerated, this approach did not allow for interstitial mPDT delivery which is desirable when treating non-superficial tumors. Hence, we have now combined the concept of interstitial fiberoptic delivery with that of a backpack light source, and report the design, fabrication and testing of two such units: one based on a diode laser and one based on an LED. Using our previous in vivo mPDT studies as a baseline, our objective was to design devices capable of delivering power levels of between 2 and 3 mW continuously over a period of 5 days [3—5]. While both of these types of sources are capable of delivering broadly comparable output powers, there are distinct advantages and drawbacks associated with each. Laser diodes offer coherent output, thus affording higher source to fiber coupling efficiencies. However, notwithstanding the incoherent nature of emitted light, the use of high luminescence LEDs as sources is advantageous from the standpoints of cost and weight. Therefore, employing devices based on either LEDs or laser diodes allows us the flexibility to perform mPDT experiments using a variety of tumor and animal models.
Materials and methods Diode laser unit For studies in tumor-bearing rabbits [6], the design is based on a diode laser-to-fiber pigtail and the prototype unit is shown in Fig. 1. A three-pin, low-power laser diode (4 mW, 635 ± 5 nm FWHM) was used as the source (Fibersense & Signals Inc., San Jose, CA). This has a threshold lasing voltage of 3.3 V at an initial current of 21 mA. The diode pins were soldered to the driver board (Fibersense & Signals Inc.), which maintained a stable (±1 mA) light output over the 5-day period by means of automatic power control (APC) of the source diode current between 20 and 22 mA. The diode with attached driver was mounted into a hollow aluminum cylin-
185
Figure 1 Tetherless, self-cooling laser diode coupled to 100 m core fiber.
der, 1.2 cm in diameter and 3.5 cm in length. The diode was pigtailed to a 100 m core optical fiber (Fibersense & Signals Inc.) using high-efficiency coupling techniques. Briefly, a multimode step index glass fiber (100 m core, 140 m outer diameter) was optically engineered to enhance its numerical aperture (NA) by forming an integral fiber lens (0.46 NA, 1 mm focal length) at its proximal end in order to maximize the capture of light from the source. The lens/fiber combination was pigtailed to the source through a specially designed adaptor head. A three-axis (xyz) translation stage was used to precisely align the source and lens/fiber combination and to vary the lens-to-source distance. The position for maximum coupling efficiency was then determined by measuring the light output at the other end of the fiber by means of an optical power meter (Newport Instruments, Newport, NJ). The fiber was then fixed into place using UV cured epoxy (Thorlabs, Newton, NJ). In earlier prototypes of this lightweight, batteryoperated device, heat generation was a significant drawback. While the increased coupling efficiency reduces this problem, the present unit also incorporates a customdesigned heat sink casing (4 cm long, 2.8 cm wide, 2.0 cm high). The anodized aluminum casing was machined with a series of thin, parallel ridges of 1 mm thickness and 3 mm spacing. In order to securely and compactly hold the battery, an internal rectangular housing (4 cm long, 0.8 cm wide, 1.7 cm high) was also machined into the casing, perpendicular to the cooling ridges. A 3.5 V, 1.6 A h Keeper II Magnum Battery (Allied Electronics) was inserted into this housing. The battery contacts were attached to the driver board via a specially designed connector cable. The total weight, including the battery’s is 42 g (diode 0.5 g, driver board 3 g, battery 16 g, heat sink, enclosure and fiber assembly 22.5 g).
LED unit For studies in small animals, units based on highluminescence LEDs were constructed. Such devices eliminate the need for a driver, and so further reduce the weight. Moreover, unlike a diode laser, an LED can readily operate over a range of voltages with little diminution in output power. This is an important consideration, given the length of the mPDT treatment, since the battery voltage inevitably drops over time.
186
Figure 2 Autocad drawing of the custom designed aluminum heat sink which features sequential, machined ridges. The LEDto-fiber pigtail is mounted into the ‘stem’ region of the T-shaped casing. The open rectangular region is has been designed to accommodate a 9.0 g, 1200 mA h, 1/2 AA lithium battery.
An LED with a luminous intensity of 8000 mcd1 and an emission wavelength of 635 nm (±5 nm, FWHM) (Super Bright LEDs Inc., St. Louis, MO) was coupled into an optical fiber of either 350 or 600 m core diameter using a custom designed integrating receptacle that includes a fiber lens and ferrule housing, via novel coupling techniques described for the laser diode-based unit. At an operating current of 15 mA, a transmitted or incident power level of 4 mW was measured. This reading was used as a reference to evaluate the efficiency of subsequent assemblies when coupling optical fibers to the diode. As with the laser diode assembly, a custom designed heat sink casing was made to house and cool the LED/fiber pigtail interface. Since there is no driver to compensate for any drop in current due to temperature rise, it is imperative that any generated heat is efficiently dissipated. This includes a heat sink area (1.9 cm × 1.7 cm × 1.0 cm) with a series of thin, machined ridges into which a hole was drilled to accommodate the diode-to-fiber interface. Adjacent to this is an integral rectangular housing (1.7 cm × 2.4 cm × 1.0 cm) to accommodate a battery holder with a prefabricated cable that allows direct connection to the LED prongs. The holder was sized to accommodate a 9.0 g, 3.3 V, 1/2 AA lithium cell battery (Keeper II Lithium PT-2150A; Allied Electronics) rated at 1200 mAh. An autocad drawing of the heat sink casing is shown in Fig. 2. The assembly was attached to the rat’s body via a rubber harness (Harvard Apparatus, Holliston, MA). The complete unit, including the harness and battery, weighs 16.5 g. Two single-fiber output device prototypes were constructed. The first employed a 600 m core fiber (pictured in Fig. 3) and the second a 350 m core fiber.
1 One thousand millicandela (mcd) equals 1 candela. This is the SI unit of luminous intensity and is the standard measure of LED output power. The candela is the luminous intensity as produced at the light source, in a given direction, of a source of monochromatic radiation of frequency 540 × 1012 Hz, with a radiant intensity in that direction of 1/683 W per steradian.
N. Davies, B.C. Wilson
Figure 3 Tetherless, high-power, self-cooling, LED-to-fibercoupled source. The high-power, single-output model featured here can deliver an initial output power of 2.4 mW at the rate of 635 nm from a 600 m core fiber. With all components in place, the device weighs 16.5 g.
Both devices were operated at room temperature over an extended period.
LED-based multifiber unit Simultaneous delivery of light through multiple optical fibers is advantageous, for example, in treating larger tumors with a uniform light dose [7]. Hence, we extended the LED-based unit by developing versions with three and five fibers coupled to the same LED source. The prototypes are shown in Figs. 4 and 5. For this, the LED was pigtailed to a fiberoptic beam splitter with three separate fiber outputs, each of 300 m core fiber. Uniform optical power splitting was achieved as follows. A 1/4 pitch lens was heat fused to the surface of the LED, thereby collimating the light from the incoherent source into a parallel beam. The output power of the LED—lens combination was 3.75 mW at an operating current of 15 mA, so that only 0.25 mW was lost. The proximal end of a single, 300 m, dual-clad glass fiber, which terminated in an engineered, integral fiber lens of NA of 0.46 (Fibersense & Signals Inc.) was placed along the horizontal axis of the LED—lens face, while the terminal end was coupled to a power meter. The proximal fiber tip was manipulated using an xyz translation stage to capture onethird of the LED output power and was then fixed in place by laser welding, in which a laser is used to fuse the fiber to the face of the lens [8—10]. The process was then repeated for the other two fibers, resulting in equal splitting within ±0.05 mW, i.e. ±4%. The entire technique was repeated
Figure 4 The initial LED/fiber pigtail design was extended to fabricate a novel version using a three-output beam splitter with 300 m core outputs. The three ports give outputs of 0.55, 0.55, and 0.65 mW, respectively allowing simultaneous irradiation of multiple tissue regions.
Tetherless fiber-coupled optical sources for extended metronomic photodynamic therapy
187
Figure 7 Optical power output as a function of time for battery-powered LED/fiber optical sources using 600 and 350 m core fiber outputs respectively.
for the five output device yielding equal splitting within ±0.025 mW, or ±2%.
meter (Photodyne, San Diego, CA) and a spectrum analyzer (Anritsu, Atsugi, Japan), respectively. The initial output was 3.2 ± 0.06 mW, and this output remained stable to within ±2% for 72 h. This corresponds to a coupling efficiency of 80%. The output power then decreased, due to the battery being unable to sustain the full current of 21 mA, reaching approximately 1 mW after 128 h. However, over the 10-day period, the emission wavelength shifted by less than 2.2 nm, demonstrating that the enhanced coupling efficiency and the cooling via the heat sink were effective. The main factor governing the variability in the output from device to device was the variation in source output between diodes, which was approximately ±5%. Other factors, such as the variations between fibers and in the fiber coupling were ±2%.
Results
LED-based units
Diode laser unit
As with the laser-based source, the LED sources were run continuously using battery power and the outputs were monitored using the spectrum analyzer. Due to the nature of LED emission, which is widely divergent and more difficult to capture using a fiber, the coupling efficiencies were lower than those obtained in the diode laser devices. The output profiles of the single-fiber units over a 4-day operation period are shown in Fig. 7. At an operating current of 15 mA, the 600 m source delivered a power of 2.4 mW, represent-
Figure 5 LED based, five fiber output optical source with equal splitting within ±0.025 mW, ±2%. This device can be readily adapted for clinical use by increasing the number and power of outputs employed.
The power output profile of this source is shown in Fig. 6. Using the Keeper II Magnum battery, the device was run continuously for 10 days, during which time, the output power and spectrum were measured using a power
Figure 6 Power output profile of the optical source powered by a single Keeper II Magnum battery. An initial power output of over 3 mW is obtained, indicating a coupling efficiency of over 80%. Higher coupling efficiencies are possible using more coherent, and costly, diodes.
Figure 8 mPDT device being worn by female Lewis rats during in vivo mPDT studies.
188 ing a coupling efficiency of over 50%. The 350 m core device delivered an initial power of 1.2 mW (coupling efficiency >25%). The output power of both sources remained constant to within ±5% for about the first 72 h. The terminal power output level of approximately 20% of the initial value was attained at day 10 (data not shown). The emission characteristics of all LED-based sources deviated by a maximum of ±4 nm from the central peak of 635 nm. This was within the tolerance range specified by the manufacturer, indicating that the heat dissipation was effective. Investigations using these LE-based sources are underway. Fig. 8 shows one device being worn by a female CNS-1 [11] tumor-bearing Lewis rat. Further experiments are underway to confirm that mPDT treatment enhances animal survival.
Discussion and conclusions These new sources fulfill two key requirements for mPDT delivery. First, they allow delivery of light via optical fiber(s), which can be stereotactically implanted to a specific depth within the tumor. Second, they are tetherless, with all components (light source, battery, circuitry, fiber and fiber coupling) contained within a self-cooling package that is sufficiently lightweight to be worn over several days by these rodent models. LED-based sources have previously been developed and used by others for the treatment of superficial tumours such as those of the skin and the oral cavity for mPDT investigations [12—14]. Nevertheless, interstitial light delivery is an important requirement for our application, given that in the established model, the tumor is induced by injecting cells to a depth of about 2 mm below the dura in order to simulate accurately the growth and PDT response of a deep-seated human tumour and to avoid the growth of the tumour outside the brain surface [15]. Light penetration through tissue is relatively poor. The optical penetration depth (OPD) is defined as the depth at which the intensity of the propagating light is attenuated to approximate 37% (1/e) of its initial value (at the air tissue interface). In brain tissue, the OPD at 635 nm is 800 m, whereas in bladder the OPD is 4 mm. Furthermore, multiple scattering within a turbid medium leads to spreading of a light beam and loss of directionality [16]. The actual light dose delivered to the tumour is, therefore, variable and difficult to determine, and this causes significant uncertainty in establishing the optimum dose—response conditions for mPDT. An impediment in prior mPDT treatments in rats has been the very poor LED-to-fiber coupling efficiency, typically in the order of a few percent [17—20]. Similarly, the previous tetherless sources employed delivered only 20—30 W from a 200 m core fiber, which is generally insufficient to overcome the tumor growth rate [21,22]. These new devices incorporate novel advances in LED-to-fiber coupling which have increased, by a factor of 60—120, the amount of light which can be delivered to the tumor. The importance of adequate, stable and continuous light delivery is emphasized by a recent study of the effectiveness of FGR-mPDT to enhance survival in VX2 bearing rabbits [5]. The mPDT source employed consisted of an LED implanted into the resection
N. Davies, B.C. Wilson cavity, and a battery pack and timer device that was held in a backpack and implanted subcutaneously. Such a device allowed the delivery of 15—30 J over 42—84 min per day or every other day. However, the animals that received combined FGR-mPDT treatment showed no significant survival advantage over animals which had received FGR alone. One possible cause for the failure of mPDT was the inability to deliver light continuously, for periods in excess of 24 h, a shortcoming that these new devices can address. In preliminary pre-clinical work, these sources have been well tolerated by rats for 4 days. A significant concern using such battery-power sources is the stability of the optical output power throughout the extended duration of the treatment. With the batteries used to date, the output has been stable to within a factor of 5% over a 3-day period. The power—time curves are different for the diode and LED sources, since the former employs a constant current driver circuit. This output stability is reasonable at this stage of development and certainly adequate to proceed to mPDT response and optimization studies. However, it is not ideal, since it does introduce another variable into these in vivo experiments and also it may be necessary to extend beyond the planned 5-day irradiation period in order to achieve ‘cure’. There are several options to ameliorate this limitation: (a) in the present design, the battery can easily be changed, without moving the optical fiber and with only a momentary period of no light; and (b) driver boards can be added that regulate the source current to extend the battery life, although this would reduce the output power and increase the weight of the device. Use of a lower capacity battery that would be changed more frequently is not a realistic solution, as the batteries used in this study offer an optimal combination of power output and weight, which is not readily provided by other commercially available battery types. A further important feature of these devices is that, in principle, the designs can be adapted for clinical use. Assuming, as results to date suggest, that the fluence rates achieved in these prototypes deliver effective mPDT in the animal model and that the PDT responses will be comparable in patients (given that we are using human tumor cell lines as part of the pre-clinical optimization), the challenge will be to deliver enough light throughout the much larger tissue volume. Hence, the total delivered power will have to be higher, probably by 1 or 2 orders of magnitude, and this will have to be spatially distributed. A possible implementation would be to implant many fibers into the brain at the time of tumor resection, since it is envisaged that mPDT would be given as an adjuvant to surgery, in the same way that acute PDT is administered at the present time for brain tumors [23]. These fibers could then be run through burr holes in the skull and under the skin to an array of lowpower LED or laser diode sources or to a higher-power single source. In turn, the sources could be either carried on the patient (as with the above tetherless units) or be externally mounted (as in the case of our earlier ‘trolley car’ design). It may even be possible to implant the whole unit under the skin, either by using high-capacity batteries or by using remote powering, as has been done, for example, for other medical devices such as pacemakers. A further significant development from standpoint of future clinical use is the highly uniform splitting ratios
Tetherless fiber-coupled optical sources for extended metronomic photodynamic therapy obtained from the multi-output devices. Wood et al. [24] reported the construction of beam-splitting devices which divided light from a PDT fiber coupled laser source into two, four and eight outputs [24]. Beam division was accomplished through the use of non-polarizing plate beam splitters, with the input power to each treatment fiber controlled with variable attenuators mounted inside the fiber couplers. Nevertheless, the sources described here obtain similar performance levels and are far less cumbersome in terms of size and construction.
Acknowledgements This work was supported by the NIH grant P01-CA43892. The authors would like to thank Annie Lin for her assistance in implanting these devices, as well as supervisory committee members Shun Wong and Lothar Lilge for helpful discussions. We also wish to extend our thanks to Fibersense & Signals Inc., San Jose, CA, for donating the lensed fibers.
References [1] Wilson BC, Patterson MS. The physics of photodynamic therapy. Phys Med Biol 1986;31:327—60. [2] Muller PJ, Wilson BC. Photodynamic therapy in primary brain tumours: clinical effects, postoperative ICP and light penetration of the brain. Photochem Photobiol 1987;46:929—35. [3] Bisland SK, et al. Metronomic photodynamic therapy as a new paradigm for the treatment of malignant brain tumors. Photochem Photobiol 2004;80:22—30. [4] Wilson BC, Bisland S, Bogaards A, et al. Metronomic photodynamic therapy (mPDT): concepts and technical feasibility in brain tumor. Proc SPIE 2003;4952:23—31. [5] Boggards A, Varma A, Zhang K, et al. Fluorescence imageguided brain tumor resection with adjuvant metronomic photodynamic therapy: pre-clinical model and technology development. Photochem Photobiol Sci 2005;4:438—42. [6] Lilge L, et al. The sensitivity of normal brain and intracranially implanted VX2 tumor to interstitial photodynamic therapy. Br J Cancer 1996;73:332—43. [7] Weersink RA, et al. Techniques for delivery and monitoring of TOOKAD (WST09)-mediated photodynamic therapy of the prostate: clinical experience and practicalities. J Photochem Photobiol B Biol 2005;79:211—22.
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[8] Karagiannis S, Chryssolouris G. Nd:YAG laser welding—–an overview. Proc SPIE 2003:5131. [9] Rathrnann FU, Russek UA. Laser welding of polymers. Using high power diode lasers. Proc SPIE 2003:5121. [10] Hoult AP. Laser welding of polymer micro-fluidic devices using novel diode laser sources. Proc SPIE 2003:5063. [11] Kruse CA, et al. A rat glioma model, CNS-1, with invasive characteristics similar to those of human gliomas: a comparison to 9L gliosarcoma. J Neurooncol 1994;22:191—200. [12] Babilas P, Kohl E, Maisch T, et al. In vitro and in vivo comparison of two different light sources for topical photodynamic therapy. Br J Dermatol 2006;154:712—8. [13] Juzeniene A, Juzenas P, Ma LW, Iani V, Moan J. Effectiveness of different light sources for 5-aminolevulinic acid photodynamic therapy. Lasers Med Sci 2004;19:139—49. [14] Taub AF. Photodynamic therapy in dermatology: history and horizons. J Drugs Dermatol 2004;3:S8—25. [15] Moriyama, et al. Bioluminescence imaging of the response of rat gliosarcoma to ALA-PpIX-mediated photodynamic therapy. Photochem Photobiol 2004;80:242—9. [16] Castano AP, Demidova TN, Hambli MR. Mechanisms in photodynamic therapy. Photodiagnos Photodyn Ther 2004;1: 279—93. [17] Berg, et al. A high performance connectorized LED package for fiber optics. IEEE Trans Components Hybrids Manufactur Technol 1981;CHMT-4:337—44. [18] Uematsu, et al. Efficient power coupling between an MH LED and a taper-ended multimode fiber. IEEE J Quant Electron 1979;QE-15:86—92. [19] Dutta, et al. High-brightness, AlGaInP-based visible lightemitting diode for efficient coupling with POF. IEEE Photon Technol Lett 1995;7:1134—6. [20] Park EH, Kim MJ, Kwon YS. Microlens for efficient coupling between LED and optical fiber. IEEE Photon Technol Lett 1999;11:439—41. [21] Hirschberg H, Spetalen S, Carper S, Hole P, Tillung T, Madsen S. Minimally invasive photodynamic therapy (PDT) for ablation of experimental rat glioma. Minim Invasive Neurosurg 2006;49:135—42. [22] Angell-Petersen E, Spetalen S, Madsen SJ, et al. Influence of light fluence rate on the effects of photodynamic therapy in an orthotopic rat glioma model. J Neurosurg 2006;104:109—17. [23] Popovic EA, Kaye AH, Hill JS. Photodynamic therapy of brain tumors. J Clin Laser Med Surg 1996;14:251—61. [24] Wood LM, Bellnier DA, Oseroff AR, Potter WR. Beam-splitting device for use with fiber-coupled laser light sources for photodynamic therapy. Photochem Photobiol 2002;76:683—5.
available at www.sciencedirect.com
journal homepage: www.elsevier.com/locate/pdpdt
Tetherless fiber-coupled optical sources for extended metronomic photodynamic therapy Nick Davies BSc (Hons), MSc a,∗, Brian C. Wilson a,b a
Department of Medical Biophysics, University of Toronto, 610 University Avenue, Toronto M5G 2M9, Canada Ontario Cancer Institute, University Health Network, 610 University Avenue, Toronto M5G 2M9, Canada Available online 8 June 2007
b
KEYWORDS Metronomic photodynamic therapy; Brain tumors; Optical sources; Light-emitting diodes; Multiport light delivery
Summary Background: Metronomic photodynamic therapy (mPDT) involves the delivery of a low fluence rate of light and photosensitizing drug, continuously and over an extended period. However, in mPDT trials, there has been a perceived need for light sources that can deliver light reliably for extended periods, while being small enough to be tolerated by small animals. Methods: We report on the development of tetherless, lightweight, fiber-coupled optical sources for in vivo delivery of interstitial mPDT. Two forms are reported, based on diode lasers and light-emitting diodes (LEDs), the latter types weighing only 16.5 g. Results: The prototypes have been well tolerated in preliminary trials on tumor-bearing rat models and can currently provide stable levels of performance for upwards of 5 days. We also report an extension of the concept to units coupled to several optical fibers to enable simultaneous irradiation of multiple locations in tissue. Conclusion: These prototypes fulfill a current need for reliable mPDT optical sources for use with animal models. They also serve as the foundation for the development of fiber-coupled sources for use in future clinical trials. © 2007 Elsevier B.V. All rights reserved.
Introduction Photodynamic therapy (PDT) is a minimally invasive form of cancer treatment that is dependent on light, a photosensitizer (PS) and oxygen in order to effect tumor cell death [1]. Conventionally, PDT is given in the form of a single, acute treatment (aPDT). In this form, the PS is delivered as a single bolus. Several hours post-drug administration, the tumor region is briefly irradiated using high-intensity light, resulting in tumor necrosis. However, aPDT treatments may also
∗
Corresponding author. Fax: +1 408 453 9293. E-mail address: [email protected] (N. Davies).
cause sensitization of normal tissue that is adjacent to the tumor and also exposed to the light field [2]. In particular, for PDT of malignant brain tumors, we have shown that damage to normal tissue, especially white matter, can be avoided and tumor cell-specific apoptotic cytotoxicity can be achieved by using aminolevulinic acid (ALA)-generated protoporphyrin IX (PpIX) if the ALA and light are delivered at low rates over an extended period [3]. This approach has been termed ‘metronomic PDT’ (mPDT) by analogy with metronomic chemotherapy, and is currently being optimized in pre-clinical rodent models bearing intracranially implanted malignant brain tumors, prior to planned clinical trials. Delivery of ALA in these animal models over an extended period is feasible by putting ALA into the drinking water, and we have shown that adequate
1572-1000/$ — see front matter © 2007 Elsevier B.V. All rights reserved. doi:10.1016/j.pdpdt.2007.03.005
Tetherless fiber-coupled optical sources for extended metronomic photodynamic therapy levels of PpIX can be achieved without significant toxicity [4]. Corresponding metronomic light delivery is more challenging, in that it is necessary to provide a means to generate and deliver the light to the intracranial tumor in a sterile, minimally invasive manner over at least several days and to administer an adequate total light fluence while also allowing the animals free movement. We have previously reported the development of both superficial and interstitial sources for mPDT delivery and the limitations associated with each [4]. In one model, the implanted optical fiber was coupled to a fixed, externally powered diode laser and the intervening length of fiber was mounted in an overhead ‘trolley car’ jig. To overcome the resulting limitations on animal mobility, we subsequently developed a battery-operated, tetherless light-emitting diode (LED) ‘backpack’. This weighed only 23 g and allowed the animal freedom of movement throughout the mPDT procedure. A long-life battery and potentiometer were housed in 5 cm × 3.5 cm × 2 cm box. The LED was placed directly into the cranium and fixed in place with bone cement over a 1 mm hole drilled in the skull. Powered by a 1.6 A h Keeper II Magnum battery (Allied Electronics, Fort Worth, TX), this device could deliver an output profile ranging from an initial power of 1.6 mW to a terminal value of 116 W over a period of 10 days. Although it was reasonably well tolerated, this approach did not allow for interstitial mPDT delivery which is desirable when treating non-superficial tumors. Hence, we have now combined the concept of interstitial fiberoptic delivery with that of a backpack light source, and report the design, fabrication and testing of two such units: one based on a diode laser and one based on an LED. Using our previous in vivo mPDT studies as a baseline, our objective was to design devices capable of delivering power levels of between 2 and 3 mW continuously over a period of 5 days [3—5]. While both of these types of sources are capable of delivering broadly comparable output powers, there are distinct advantages and drawbacks associated with each. Laser diodes offer coherent output, thus affording higher source to fiber coupling efficiencies. However, notwithstanding the incoherent nature of emitted light, the use of high luminescence LEDs as sources is advantageous from the standpoints of cost and weight. Therefore, employing devices based on either LEDs or laser diodes allows us the flexibility to perform mPDT experiments using a variety of tumor and animal models.
Materials and methods Diode laser unit For studies in tumor-bearing rabbits [6], the design is based on a diode laser-to-fiber pigtail and the prototype unit is shown in Fig. 1. A three-pin, low-power laser diode (4 mW, 635 ± 5 nm FWHM) was used as the source (Fibersense & Signals Inc., San Jose, CA). This has a threshold lasing voltage of 3.3 V at an initial current of 21 mA. The diode pins were soldered to the driver board (Fibersense & Signals Inc.), which maintained a stable (±1 mA) light output over the 5-day period by means of automatic power control (APC) of the source diode current between 20 and 22 mA. The diode with attached driver was mounted into a hollow aluminum cylin-
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Figure 1 Tetherless, self-cooling laser diode coupled to 100 m core fiber.
der, 1.2 cm in diameter and 3.5 cm in length. The diode was pigtailed to a 100 m core optical fiber (Fibersense & Signals Inc.) using high-efficiency coupling techniques. Briefly, a multimode step index glass fiber (100 m core, 140 m outer diameter) was optically engineered to enhance its numerical aperture (NA) by forming an integral fiber lens (0.46 NA, 1 mm focal length) at its proximal end in order to maximize the capture of light from the source. The lens/fiber combination was pigtailed to the source through a specially designed adaptor head. A three-axis (xyz) translation stage was used to precisely align the source and lens/fiber combination and to vary the lens-to-source distance. The position for maximum coupling efficiency was then determined by measuring the light output at the other end of the fiber by means of an optical power meter (Newport Instruments, Newport, NJ). The fiber was then fixed into place using UV cured epoxy (Thorlabs, Newton, NJ). In earlier prototypes of this lightweight, batteryoperated device, heat generation was a significant drawback. While the increased coupling efficiency reduces this problem, the present unit also incorporates a customdesigned heat sink casing (4 cm long, 2.8 cm wide, 2.0 cm high). The anodized aluminum casing was machined with a series of thin, parallel ridges of 1 mm thickness and 3 mm spacing. In order to securely and compactly hold the battery, an internal rectangular housing (4 cm long, 0.8 cm wide, 1.7 cm high) was also machined into the casing, perpendicular to the cooling ridges. A 3.5 V, 1.6 A h Keeper II Magnum Battery (Allied Electronics) was inserted into this housing. The battery contacts were attached to the driver board via a specially designed connector cable. The total weight, including the battery’s is 42 g (diode 0.5 g, driver board 3 g, battery 16 g, heat sink, enclosure and fiber assembly 22.5 g).
LED unit For studies in small animals, units based on highluminescence LEDs were constructed. Such devices eliminate the need for a driver, and so further reduce the weight. Moreover, unlike a diode laser, an LED can readily operate over a range of voltages with little diminution in output power. This is an important consideration, given the length of the mPDT treatment, since the battery voltage inevitably drops over time.
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Figure 2 Autocad drawing of the custom designed aluminum heat sink which features sequential, machined ridges. The LEDto-fiber pigtail is mounted into the ‘stem’ region of the T-shaped casing. The open rectangular region is has been designed to accommodate a 9.0 g, 1200 mA h, 1/2 AA lithium battery.
An LED with a luminous intensity of 8000 mcd1 and an emission wavelength of 635 nm (±5 nm, FWHM) (Super Bright LEDs Inc., St. Louis, MO) was coupled into an optical fiber of either 350 or 600 m core diameter using a custom designed integrating receptacle that includes a fiber lens and ferrule housing, via novel coupling techniques described for the laser diode-based unit. At an operating current of 15 mA, a transmitted or incident power level of 4 mW was measured. This reading was used as a reference to evaluate the efficiency of subsequent assemblies when coupling optical fibers to the diode. As with the laser diode assembly, a custom designed heat sink casing was made to house and cool the LED/fiber pigtail interface. Since there is no driver to compensate for any drop in current due to temperature rise, it is imperative that any generated heat is efficiently dissipated. This includes a heat sink area (1.9 cm × 1.7 cm × 1.0 cm) with a series of thin, machined ridges into which a hole was drilled to accommodate the diode-to-fiber interface. Adjacent to this is an integral rectangular housing (1.7 cm × 2.4 cm × 1.0 cm) to accommodate a battery holder with a prefabricated cable that allows direct connection to the LED prongs. The holder was sized to accommodate a 9.0 g, 3.3 V, 1/2 AA lithium cell battery (Keeper II Lithium PT-2150A; Allied Electronics) rated at 1200 mAh. An autocad drawing of the heat sink casing is shown in Fig. 2. The assembly was attached to the rat’s body via a rubber harness (Harvard Apparatus, Holliston, MA). The complete unit, including the harness and battery, weighs 16.5 g. Two single-fiber output device prototypes were constructed. The first employed a 600 m core fiber (pictured in Fig. 3) and the second a 350 m core fiber.
1 One thousand millicandela (mcd) equals 1 candela. This is the SI unit of luminous intensity and is the standard measure of LED output power. The candela is the luminous intensity as produced at the light source, in a given direction, of a source of monochromatic radiation of frequency 540 × 1012 Hz, with a radiant intensity in that direction of 1/683 W per steradian.
N. Davies, B.C. Wilson
Figure 3 Tetherless, high-power, self-cooling, LED-to-fibercoupled source. The high-power, single-output model featured here can deliver an initial output power of 2.4 mW at the rate of 635 nm from a 600 m core fiber. With all components in place, the device weighs 16.5 g.
Both devices were operated at room temperature over an extended period.
LED-based multifiber unit Simultaneous delivery of light through multiple optical fibers is advantageous, for example, in treating larger tumors with a uniform light dose [7]. Hence, we extended the LED-based unit by developing versions with three and five fibers coupled to the same LED source. The prototypes are shown in Figs. 4 and 5. For this, the LED was pigtailed to a fiberoptic beam splitter with three separate fiber outputs, each of 300 m core fiber. Uniform optical power splitting was achieved as follows. A 1/4 pitch lens was heat fused to the surface of the LED, thereby collimating the light from the incoherent source into a parallel beam. The output power of the LED—lens combination was 3.75 mW at an operating current of 15 mA, so that only 0.25 mW was lost. The proximal end of a single, 300 m, dual-clad glass fiber, which terminated in an engineered, integral fiber lens of NA of 0.46 (Fibersense & Signals Inc.) was placed along the horizontal axis of the LED—lens face, while the terminal end was coupled to a power meter. The proximal fiber tip was manipulated using an xyz translation stage to capture onethird of the LED output power and was then fixed in place by laser welding, in which a laser is used to fuse the fiber to the face of the lens [8—10]. The process was then repeated for the other two fibers, resulting in equal splitting within ±0.05 mW, i.e. ±4%. The entire technique was repeated
Figure 4 The initial LED/fiber pigtail design was extended to fabricate a novel version using a three-output beam splitter with 300 m core outputs. The three ports give outputs of 0.55, 0.55, and 0.65 mW, respectively allowing simultaneous irradiation of multiple tissue regions.
Tetherless fiber-coupled optical sources for extended metronomic photodynamic therapy
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Figure 7 Optical power output as a function of time for battery-powered LED/fiber optical sources using 600 and 350 m core fiber outputs respectively.
for the five output device yielding equal splitting within ±0.025 mW, or ±2%.
meter (Photodyne, San Diego, CA) and a spectrum analyzer (Anritsu, Atsugi, Japan), respectively. The initial output was 3.2 ± 0.06 mW, and this output remained stable to within ±2% for 72 h. This corresponds to a coupling efficiency of 80%. The output power then decreased, due to the battery being unable to sustain the full current of 21 mA, reaching approximately 1 mW after 128 h. However, over the 10-day period, the emission wavelength shifted by less than 2.2 nm, demonstrating that the enhanced coupling efficiency and the cooling via the heat sink were effective. The main factor governing the variability in the output from device to device was the variation in source output between diodes, which was approximately ±5%. Other factors, such as the variations between fibers and in the fiber coupling were ±2%.
Results
LED-based units
Diode laser unit
As with the laser-based source, the LED sources were run continuously using battery power and the outputs were monitored using the spectrum analyzer. Due to the nature of LED emission, which is widely divergent and more difficult to capture using a fiber, the coupling efficiencies were lower than those obtained in the diode laser devices. The output profiles of the single-fiber units over a 4-day operation period are shown in Fig. 7. At an operating current of 15 mA, the 600 m source delivered a power of 2.4 mW, represent-
Figure 5 LED based, five fiber output optical source with equal splitting within ±0.025 mW, ±2%. This device can be readily adapted for clinical use by increasing the number and power of outputs employed.
The power output profile of this source is shown in Fig. 6. Using the Keeper II Magnum battery, the device was run continuously for 10 days, during which time, the output power and spectrum were measured using a power
Figure 6 Power output profile of the optical source powered by a single Keeper II Magnum battery. An initial power output of over 3 mW is obtained, indicating a coupling efficiency of over 80%. Higher coupling efficiencies are possible using more coherent, and costly, diodes.
Figure 8 mPDT device being worn by female Lewis rats during in vivo mPDT studies.
188 ing a coupling efficiency of over 50%. The 350 m core device delivered an initial power of 1.2 mW (coupling efficiency >25%). The output power of both sources remained constant to within ±5% for about the first 72 h. The terminal power output level of approximately 20% of the initial value was attained at day 10 (data not shown). The emission characteristics of all LED-based sources deviated by a maximum of ±4 nm from the central peak of 635 nm. This was within the tolerance range specified by the manufacturer, indicating that the heat dissipation was effective. Investigations using these LE-based sources are underway. Fig. 8 shows one device being worn by a female CNS-1 [11] tumor-bearing Lewis rat. Further experiments are underway to confirm that mPDT treatment enhances animal survival.
Discussion and conclusions These new sources fulfill two key requirements for mPDT delivery. First, they allow delivery of light via optical fiber(s), which can be stereotactically implanted to a specific depth within the tumor. Second, they are tetherless, with all components (light source, battery, circuitry, fiber and fiber coupling) contained within a self-cooling package that is sufficiently lightweight to be worn over several days by these rodent models. LED-based sources have previously been developed and used by others for the treatment of superficial tumours such as those of the skin and the oral cavity for mPDT investigations [12—14]. Nevertheless, interstitial light delivery is an important requirement for our application, given that in the established model, the tumor is induced by injecting cells to a depth of about 2 mm below the dura in order to simulate accurately the growth and PDT response of a deep-seated human tumour and to avoid the growth of the tumour outside the brain surface [15]. Light penetration through tissue is relatively poor. The optical penetration depth (OPD) is defined as the depth at which the intensity of the propagating light is attenuated to approximate 37% (1/e) of its initial value (at the air tissue interface). In brain tissue, the OPD at 635 nm is 800 m, whereas in bladder the OPD is 4 mm. Furthermore, multiple scattering within a turbid medium leads to spreading of a light beam and loss of directionality [16]. The actual light dose delivered to the tumour is, therefore, variable and difficult to determine, and this causes significant uncertainty in establishing the optimum dose—response conditions for mPDT. An impediment in prior mPDT treatments in rats has been the very poor LED-to-fiber coupling efficiency, typically in the order of a few percent [17—20]. Similarly, the previous tetherless sources employed delivered only 20—30 W from a 200 m core fiber, which is generally insufficient to overcome the tumor growth rate [21,22]. These new devices incorporate novel advances in LED-to-fiber coupling which have increased, by a factor of 60—120, the amount of light which can be delivered to the tumor. The importance of adequate, stable and continuous light delivery is emphasized by a recent study of the effectiveness of FGR-mPDT to enhance survival in VX2 bearing rabbits [5]. The mPDT source employed consisted of an LED implanted into the resection
N. Davies, B.C. Wilson cavity, and a battery pack and timer device that was held in a backpack and implanted subcutaneously. Such a device allowed the delivery of 15—30 J over 42—84 min per day or every other day. However, the animals that received combined FGR-mPDT treatment showed no significant survival advantage over animals which had received FGR alone. One possible cause for the failure of mPDT was the inability to deliver light continuously, for periods in excess of 24 h, a shortcoming that these new devices can address. In preliminary pre-clinical work, these sources have been well tolerated by rats for 4 days. A significant concern using such battery-power sources is the stability of the optical output power throughout the extended duration of the treatment. With the batteries used to date, the output has been stable to within a factor of 5% over a 3-day period. The power—time curves are different for the diode and LED sources, since the former employs a constant current driver circuit. This output stability is reasonable at this stage of development and certainly adequate to proceed to mPDT response and optimization studies. However, it is not ideal, since it does introduce another variable into these in vivo experiments and also it may be necessary to extend beyond the planned 5-day irradiation period in order to achieve ‘cure’. There are several options to ameliorate this limitation: (a) in the present design, the battery can easily be changed, without moving the optical fiber and with only a momentary period of no light; and (b) driver boards can be added that regulate the source current to extend the battery life, although this would reduce the output power and increase the weight of the device. Use of a lower capacity battery that would be changed more frequently is not a realistic solution, as the batteries used in this study offer an optimal combination of power output and weight, which is not readily provided by other commercially available battery types. A further important feature of these devices is that, in principle, the designs can be adapted for clinical use. Assuming, as results to date suggest, that the fluence rates achieved in these prototypes deliver effective mPDT in the animal model and that the PDT responses will be comparable in patients (given that we are using human tumor cell lines as part of the pre-clinical optimization), the challenge will be to deliver enough light throughout the much larger tissue volume. Hence, the total delivered power will have to be higher, probably by 1 or 2 orders of magnitude, and this will have to be spatially distributed. A possible implementation would be to implant many fibers into the brain at the time of tumor resection, since it is envisaged that mPDT would be given as an adjuvant to surgery, in the same way that acute PDT is administered at the present time for brain tumors [23]. These fibers could then be run through burr holes in the skull and under the skin to an array of lowpower LED or laser diode sources or to a higher-power single source. In turn, the sources could be either carried on the patient (as with the above tetherless units) or be externally mounted (as in the case of our earlier ‘trolley car’ design). It may even be possible to implant the whole unit under the skin, either by using high-capacity batteries or by using remote powering, as has been done, for example, for other medical devices such as pacemakers. A further significant development from standpoint of future clinical use is the highly uniform splitting ratios
Tetherless fiber-coupled optical sources for extended metronomic photodynamic therapy obtained from the multi-output devices. Wood et al. [24] reported the construction of beam-splitting devices which divided light from a PDT fiber coupled laser source into two, four and eight outputs [24]. Beam division was accomplished through the use of non-polarizing plate beam splitters, with the input power to each treatment fiber controlled with variable attenuators mounted inside the fiber couplers. Nevertheless, the sources described here obtain similar performance levels and are far less cumbersome in terms of size and construction.
Acknowledgements This work was supported by the NIH grant P01-CA43892. The authors would like to thank Annie Lin for her assistance in implanting these devices, as well as supervisory committee members Shun Wong and Lothar Lilge for helpful discussions. We also wish to extend our thanks to Fibersense & Signals Inc., San Jose, CA, for donating the lensed fibers.
References [1] Wilson BC, Patterson MS. The physics of photodynamic therapy. Phys Med Biol 1986;31:327—60. [2] Muller PJ, Wilson BC. Photodynamic therapy in primary brain tumours: clinical effects, postoperative ICP and light penetration of the brain. Photochem Photobiol 1987;46:929—35. [3] Bisland SK, et al. Metronomic photodynamic therapy as a new paradigm for the treatment of malignant brain tumors. Photochem Photobiol 2004;80:22—30. [4] Wilson BC, Bisland S, Bogaards A, et al. Metronomic photodynamic therapy (mPDT): concepts and technical feasibility in brain tumor. Proc SPIE 2003;4952:23—31. [5] Boggards A, Varma A, Zhang K, et al. Fluorescence imageguided brain tumor resection with adjuvant metronomic photodynamic therapy: pre-clinical model and technology development. Photochem Photobiol Sci 2005;4:438—42. [6] Lilge L, et al. The sensitivity of normal brain and intracranially implanted VX2 tumor to interstitial photodynamic therapy. Br J Cancer 1996;73:332—43. [7] Weersink RA, et al. Techniques for delivery and monitoring of TOOKAD (WST09)-mediated photodynamic therapy of the prostate: clinical experience and practicalities. J Photochem Photobiol B Biol 2005;79:211—22.
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[8] Karagiannis S, Chryssolouris G. Nd:YAG laser welding—–an overview. Proc SPIE 2003:5131. [9] Rathrnann FU, Russek UA. Laser welding of polymers. Using high power diode lasers. Proc SPIE 2003:5121. [10] Hoult AP. Laser welding of polymer micro-fluidic devices using novel diode laser sources. Proc SPIE 2003:5063. [11] Kruse CA, et al. A rat glioma model, CNS-1, with invasive characteristics similar to those of human gliomas: a comparison to 9L gliosarcoma. J Neurooncol 1994;22:191—200. [12] Babilas P, Kohl E, Maisch T, et al. In vitro and in vivo comparison of two different light sources for topical photodynamic therapy. Br J Dermatol 2006;154:712—8. [13] Juzeniene A, Juzenas P, Ma LW, Iani V, Moan J. Effectiveness of different light sources for 5-aminolevulinic acid photodynamic therapy. Lasers Med Sci 2004;19:139—49. [14] Taub AF. Photodynamic therapy in dermatology: history and horizons. J Drugs Dermatol 2004;3:S8—25. [15] Moriyama, et al. Bioluminescence imaging of the response of rat gliosarcoma to ALA-PpIX-mediated photodynamic therapy. Photochem Photobiol 2004;80:242—9. [16] Castano AP, Demidova TN, Hambli MR. Mechanisms in photodynamic therapy. Photodiagnos Photodyn Ther 2004;1: 279—93. [17] Berg, et al. A high performance connectorized LED package for fiber optics. IEEE Trans Components Hybrids Manufactur Technol 1981;CHMT-4:337—44. [18] Uematsu, et al. Efficient power coupling between an MH LED and a taper-ended multimode fiber. IEEE J Quant Electron 1979;QE-15:86—92. [19] Dutta, et al. High-brightness, AlGaInP-based visible lightemitting diode for efficient coupling with POF. IEEE Photon Technol Lett 1995;7:1134—6. [20] Park EH, Kim MJ, Kwon YS. Microlens for efficient coupling between LED and optical fiber. IEEE Photon Technol Lett 1999;11:439—41. [21] Hirschberg H, Spetalen S, Carper S, Hole P, Tillung T, Madsen S. Minimally invasive photodynamic therapy (PDT) for ablation of experimental rat glioma. Minim Invasive Neurosurg 2006;49:135—42. [22] Angell-Petersen E, Spetalen S, Madsen SJ, et al. Influence of light fluence rate on the effects of photodynamic therapy in an orthotopic rat glioma model. J Neurosurg 2006;104:109—17. [23] Popovic EA, Kaye AH, Hill JS. Photodynamic therapy of brain tumors. J Clin Laser Med Surg 1996;14:251—61. [24] Wood LM, Bellnier DA, Oseroff AR, Potter WR. Beam-splitting device for use with fiber-coupled laser light sources for photodynamic therapy. Photochem Photobiol 2002;76:683—5.
