Abstract:
A method and apparatus for providing a miniaturized, flexible high voltage up-converter. Aspects of the invention are particularly useful in providing an apparatus comprising a plurality of up-converting modules while also allowing the apparatus to maintain a desired degree of flexibility. However, certain aspects of the invention may be equally applicable in other scenarios as well.

Description:
BACKGROUND 
   I. Field of the Invention 
   The present invention is generally directed to a method and apparatus for voltage upconverting. More particularly, the present invention is directed to a method and apparatus for providing a miniaturized, flexible high voltage up-converter. Aspects of the invention are particularly useful in providing an apparatus comprising a plurality of up-converting modules while also allowing the apparatus to maintain a desired degree of flexibility. However, certain aspects of the invention may be equally applicable in other scenarios as well. 
   II. Description of Related Technology 
   In conventional angioplasty operations, a stent is inserted into a patient&#39;s artery that may be occluded or constricted by plaque. These stents allow a surgeon to, via in-vivo stent manipulation and guidance, enter the patient&#39;s body and keep the occlusion unrestricted. If, subsequently, the stent occluded again, irradiation of the constricted stent area may be required. Alternatively, catheters may be used to irradiate a cancerous growth. Irradiation occurs with radioactive seeds emitting beta or gamma rays. To produce these beta or gamma rays, however, the catheters have to be provided with means allowing those radioactive seeds to travel to the site where treatment is needed. These seeds are highly radioactive and irradiate the entire length of the artery from insertion point to the treatment site. Once removed from the shielding container, they pose a health risk to the patient and the medical professionals administering treatment. To solve some of these problems associated with the radioactive seeds, it is desirable to provide an X-ray generating device that generates an X-ray source near the desired area and the X-ray dose can be generated in-vivo at will. Because the radiation must be produced at the site of interest (i.e., at the obstructed artery or the cancerous growth), this X-ray tube is typically located at a distal end of the stent or catheter. 
   Under ordinary operation, these X-ray tubes require a high degree of power (voltage) to operate. For example, U.S. Pat. No. 5,090,043 entitled “X-ray Micro-Tube and Method of Use in Radiation Oncology” to Parker et al. teaches the use of an apparatus and method for the treatment of a patient having a tumor. 
   Parker et al. teaches using an X-ray generating source positioned at a location in close proximity to site or application (e.g., an artery, a vein, or a tumor). The X-ray generating source is operable at a voltage level in the range of approximately 10-60 kilo-electron volts (keV) to thereby enhance absorption of the generated X-rays by the tumor and minimizing the side effects of radiation therapy on the patient normal tissue. 
   Therefore, to provide the necessary voltage and power to certain types of miniature X-ray tubes, approximately 10-60 keV are required. For treatment of occluded stents, approximately 20 kV are sufficient. To provide this high level of voltage at a distal end of a catheter, the power must be provided along the entire catheter cable to the catheter distal end. Proposed catheters, however, are provided with a lengthy high voltage power cable. For example, a proposed catheter high voltage cable is typically on the order of three feet in length. Therefore, a dangerous situation arises where the peak voltage of 20 kV must traverse along the entire length of the catheter cable and then along the length of the catheter to eventually reach the X-ray unit. 
   Providing this peak voltage along a high voltage cable feeding the catheter and then also running the entire length of the catheter poses certain dangerous operating conditions. For example, storing such a large amount of energy can accidentally and/or inadvertently discharge and harm or fatally injure a patient and/or physician. Flashover between the high voltage components and an exterior housing of the catheter (an electrical ground) could harm or even kill the patient, the administering physician, and/or others involved in the in-vivo operation (e.g., members of the operating staff). Flashover occurs where there is leakage between the outside grounded and the inside high-voltage and this leakage if followed by a dielectric breakdown. This flashover concern exists along the entire cable length. 
   Therefore, because of a requirement for a long high voltage cable, most proposed miniature X-ray tube catheter systems behave as essentially a very large, charged capacitor. 
   There is, therefore, a general need to be able to reduce the necessity of a lengthy high voltage cable. There is also a general need to reduce a high voltage system&#39;s overall capacitance, and therefore potential flashover. These general needs should also be met while also being able to generate a high enough voltage for X-ray application at the point of observation or X-ray application. 
   Aside from these high voltage breakdown and capacitance concerns, there is also a maneuvering or manipulating concern associated with catheters containing miniaturized X-ray tubes. For example, because such medical devices may be used in a variety of applications (e.g., angioplasty, tumor irradiation, etc.), the catheter containing certain components must be flexible enough so that during an in-vivo operation, a user of the device (i.e., a surgeon) can maneuver and/or manipulate the subject catheter so as to manipulate or guide the X-ray unit along an artery to accurately position the catheter at a desired location. Because certain proposed miniature X-ray tubes have been large (larger than 2.5 mm in diameter), there is a further need for a flexible, guidable device comprising a X-ray tube device having a diameter less than 2.5 mm. It is believed that an ideal X-ray tube for angioplasty procedures has a diameter ranging from 0.5 to 1.0 mm. Other diameter sizes could also be desired depending on the application of the X-ray tube. 
   Even though the same proposed concepts may describe a relatively compact voltage source, the dimensions of such voltage sources are large, often on the scale of inches. For example, the voltage source disclosed in U.S. Pat. No. 4,241,360 discloses a voltage source having a size on the order of 0.5 inches in length (12 mm) and 0.2 inches wide (5 mm). For medical applications, and especially for catheters that are inserted inside a body, further miniaturization is desired. In addition, where a voltage source is used in a catheter or other in-vivo applications, the catheter and hence the voltage source must have some degree of flexibility and maneuverability. 
   There is also a need to provide a miniaturized power source that reduces the risk to the patient by minimizing the discharge power and while also maintaining catheter flexibility. There is also a general need to provide a power source that can be guidable through difficult passageways and provide a source of power at difficult to reach areas. 
   SUMMARY 
   According to one exemplary arrangement, an apparatus for up-converting an initial voltage includes a first photodiode module. A first photon source is optically coupled to the first photodiode module to a light source. The first module up-converts the initial voltage to a first up-converted voltage. A second photodiode module receives the first up-converted voltage from the first photodiode module and a photon source coupled to the second photodiode module to the source. The second photodiode module up-converts the first up-converted voltage to a second up-converted voltage. 
   According to another exemplary arrangement, a method of voltage up-converting includes the steps of electrically coupling a first up-converting module to a second module and optically coupling a photon source to the first and the second modules. The photon source generates at the first up-converting module an up-converted voltage. This upconverted voltage is provided to the second up-converting module from the first up-converting module and the source is utilized to up-convert the first voltage. 
   In another arrangement, a miniaturized, flexible voltage up-converting instrument includes an X-ray generating source insertable into a body of a patient to a location in close proximity to a desired point of X-ray application. A first modular voltage up-converter is coupled to the generating source. An up-converted voltage is applied to the source to generate a desired X-ray dose at a desired point of application. 
   In yet an alternative arrangement, a method of fabricating a voltage up-converter is provided. The method includes the steps of:
     providing a substrate containing a thin layer of silicon conducting material;   performing an initial n −  type of ion implantation region, said region defining an initial photodiode pattern on a surface of said substrate;   fabricating a plurality of p +  type ion implantation regions along said photodiode pattern;   fabricating a second plurality of n +  ion implantation regions along said photodiode pattern; and   providing an oxide layer along a top surface of said photodiode pattern.   

   In yet another arrangement, a miniaturized, flexible voltage up-converting instrument includes an X-ray generating source insertable into a body of a patient to a location in close proximity to a desired point of X-ray application. A first modular voltage up-converter is coupled to the X-ray generating source so that an up-converted voltage is applied to the X-ray source to generate a desired amount of X-ray dose at the desired point of X-ray application. 
   These as well as other advantages of various aspects of applicant&#39;s present arrangements will become apparent to those of ordinary skill in the art by reading the following detailed description, with appropriate reference to the accompanying drawings. 

   
     BRIEF DESCRIPTION OF THE DRAWINGS 
     An exemplary arrangement described herein with reference to the drawings, in which: 
       FIG. 1  illustrates a known miniature X-ray apparatus system; 
       FIG. 2  is a cross sectional view of the high-voltage cable of a miniature X-ray apparatus system illustrated in  FIG. 1 ; 
       FIG. 3  illustrates a schematic representation of a preferred arrangement of a miniature, flexible voltage up-converter system comprising a plurality of voltage up-converter modules; 
       FIG. 4  illustrates a schematic representation of one of the plurality of voltage up-converter modules provided in the up-converter system illustrated in  FIG. 3 ; 
       FIG. 5  illustrates a flow chart providing certain processing steps for fabricating an arrangement of the voltage up-converter illustrated in  FIG. 4 ; 
     FIG.  6 ( a ) illustrates a first processing step for fabricating a lateral type photodiode pattern that can be used for the voltage up-converter module illustrated in  FIG. 4 ; 
     FIG.  6 ( b ) illustrates another processing step for fabricating the voltage up-converter illustrated in  FIG. 4 ; 
     FIG.  6 ( c ) illustrates another processing step for fabricating the voltage up-converter illustrated in  FIG. 4 ; 
     FIG.  6 ( d ) illustrates another processing step for fabricating the voltage up-converter illustrated in  FIG. 4 ; 
     FIG.  6 ( e ) illustrates another processing step for fabricating the voltage up-converter illustrated in  FIG. 4 ; 
     FIG.  6 ( f ) illustrates another processing step for fabricating the voltage up-converter illustrated in  FIG. 4 ; 
     FIG.  6 ( g ) illustrates a cross sectional view of the voltage up-converter illustrated in FIG.  6 ( f ); 
     FIG.  6 ( h ) illustrates another processing step for fabricating the voltage up-converter illustrated in  FIG. 4 ; 
     FIG.  6 ( i ) illustrates a cross section of the FIG.  6 ( h ) including an additional processing step; 
       FIG. 7  illustrates a representative voltage (V) versus current (I) graph of one of the photo-diodes included in the voltage up-converter section illustrated in FIG.  6 ( a ); 
       FIG. 8  illustrates a perspective view of the voltage up-converter module illustrated in  FIG. 3  fabricated on a sapphire substrate; 
       FIG. 9  is a side view of a voltage up-converter illustrated in  FIG. 8  coupled to a light source; 
     FIG.  10 ( a ) illustrates a preferred arrangement of an encapsulated voltage-up-converter module; and 
     FIG.  10 ( b ) illustrates another preferred arrangement of an encapsulated voltage-up-converter module. 
   

   DETAILED DESCRIPTION 
   1. Overview 
   As previously described, X-ray generating devices require a large potential voltage. For example, a proposed X-ray device, such as the X-ray device illustrated in  FIG. 1 , generally requires an applied voltage on the order of between 15 kV to 30 kV. Providing such a large potential voltage presents certain safety concerns, especially where the X-ray generator is provided in a miniaturized instrument, such as a catheter. For example, one such typical X-ray device system is illustrated in FIG.  1 . 
     FIG. 1  illustrates a schematic view of a proposed arrangement of a high voltage X-ray system  10 . The X-ray system  10  includes an instrumentation system  18  and an X-ray enclosure  14 . The enclosure  14  contains an X-ray emitting apparatus  12 . Apparatus  12  includes an X-ray emitting source and a high voltage wire  16 . The X-ray emitting source is located at a distal end of the enclosure. The X-ray emitting source  12  must be electrically coupled to the high voltage source  22 , such as by way of a high voltage wire  16 . 
   As seen from  FIG. 1 , the X-ray source  12  includes an X-ray emitting head  21  and a power wire  16  to which the head  21  is connected. The instrumentation system  18  is also provided and includes a control unit  20  and a high voltage power source  22 . The control unit  20 , preferably an operator controlled unit, operates the X-ray unit and determines, via an operator control device, when the X-ray apparatus begins irradiation. The operator control device (not shown) could be a foot switch or other human interface, such as a button, switch, or other like device. 
   As illustrated in  FIG. 1 , the X-ray apparatus is directly coupled to the high voltage wire  16 . The high voltage cable runs the length of the X-ray system housing  14 , L h . At one end d of the X-ray system housing  14 , the high voltage wire terminates at the instrumentation system  18  where the high voltage wire is electrically coupled to the high voltage power source  22 . 
   Typically, the X-ray head  21  will include a vacuum chamber. The vacuum chamber houses a microscopic cathode for generating electrons. An anode will also be provided. The anode accelerates and attracts the electrons and emits X-ray radiation  24  upon bombardment by the accelerated electrons. The emitted X-ray radiation is then used to irradiate a constricted artery, a cancerous growth (tumor), or other unwanted substance. For more information relating to such a typical X-ray head, the reader is directed to Tang U.S. Pat. No. 5,729,583; Parker U.S. Pat. No. 5,090,043; Smith U.S. Pat. No. 5,984,853; and Smith U.S. Pat. No. 6,241,651, herein entirely incorporated by reference and to which the reader is directed for further details. 
   As illustrated, the X-ray emitter  12  resides within an enclosure  14 . Such an enclosure could include a manually manipulated medical device used for in-vivo applications. For example, the housing illustrated in  FIG. 1  could be a catheter used in an angioplasty operation. Alternatively, the miniaturized X-ray source could be used or placed within the confines of a structure that requires a high potential including dental applications, desk top crystallography, protein examinations, and the like. 
   One limitation as to the manipulation of the arrangement of  FIG. 1  relates to the actual size of the X-ray unit  12  (L x ) and its diameter and the high voltage wire provided along the length L h  of the body. Where the X-ray unit  12  has a length of L x , the catheter head  21  would not be able to be flexed along this portion of the enclosure. Rather, the housing could only be flexed in between the points of b and d and could not be flexed between the points of c and b. The actual size of the catheter head  21  therefore, restrains the manipulation of the enclosure. 
   Another limitation of the arrangement illustrated in  FIG. 1  stems from the fact that the high voltage wire  16  carries the high voltage the entire distance L h  from the power source to the X-ray unit  12 . Therefore, there is a potential risk that there will be a breakdown between the internal wire and the catheter outer enclosure  26  (i.e., ground). This is particularly problematic given that a potential breakdown could occur anywhere along the entire length of the catheter since a voltage of significant magnitude is present along the entire length of the cable. This is particularly problematic given that the enclosure may be used as an in-vivo medical instrument. 
   The X-ray apparatus  14  comprises both a distal end and a proximal end. Where the high voltage power source is utilized to generate X-rays, the high voltage power source will ordinarily be located in the proximal end of the medical device. Once the X-ray unit is energized with a desired amount of power, X-rays  24  are emitted from the distal end. Preferably, these X-rays  24  are emitted in a rotational symmetric fashion. 
     FIG. 2  illustrates a cross sectional view of the device housing illustrated in  FIG. 1  along the A-A′ view. Referring now to  FIGS. 1 and 2 , typically, the enclosure housing  26  of the system  10  has a diameter d ranging from about 1.4 to about 2 millimeter (mm). Diameter d may be measured from the center of the high voltage wire  16  to the outer wall  28  of enclosure  26 . A proposed X-ray device, such as the X-ray device  12  illustrated in  FIG. 1 , may have a diameter of approximately 1-2 mm and may have a length L x  of approximately 2-4 mm long. The catheter extends from the proximal end to the distal end wherein this length L h  could be as long as 3 feet. Supplying a high voltage (20 kV) along a 3 foot cable having a diameter of about 2 mm presents a dangerous situation since the energy stored along the catheter is large and therefore, the catheter in essence acts like a large, charged capacitor. 
   This may be seen by equating the energy stored in such a system. For example, the stored energy of the system may be calculated using the following equation: U=½C*V 2 , where U is the stored potential energy, C is the capacitance, and V the voltage. For a catheter having a diameter of approximately 1.4 mm and the high voltage wire having a radius r 1  ranging from 40 to approximately 220 micrometers (μm), the overall capacitance of the device will generally range from 3.9×10 −11  to 9.6×10 −11  Farads for a 3 foot cable and catheter system. If the X-ray source  12  required about 20 kV of power, the total energy stored along the device would approach 0.01 Watt-seconds. If this energy were to inadvertently discharge during a short time period, for example during 1 micro-second (μsec) time interval, the dissipated power (P=U/T) would be at a dangerous level: P=10 4  Watts. Therefore, by providing a point of use power supply (providing a desired amount of power at only one specific point), which is in close proximity to the X-ray head, the overall system capacitance may be significantly reduced and therefore the capacitive discharge potential. 
     FIG. 3  illustrates one preferred arrangement of a voltage up-converter system  50 . The system  50  includes a voltage up-converting point of use device  52  coupled to an instrumentation control  54 . The instrumentation control  54  includes a power source  56 , a light source  58 , and a control circuit  57 . 
   The up-converting point of use device  52  extends from a distal end  51  to a proximal end  53  and includes a plurality of voltage up-converter modules  62 ( a-d ). An X-ray emitter  64  is provided at the distal end. As illustrated in  FIG. 3 , the up-converting point of use device includes four up-converter modules are shown. However, it will be appreciated by those of ordinary skill in the art that other up-converter modules may also be utilized. For example, a point of use device could have more or less than four modules depending on the overall design and performance requirements sought. 
   As will be described in further detail below, each voltage up-converter module  62 ( a-d ) comprises a plurality of photodiodes. As can be seen from the arrangement illustrated in  FIG. 3 , the modules  62 ( a-d ) are coupled in a cascaded series, one after the other. Alternatively, a miniaturized up-converting module could comprise, rather than photodiodes, certain conventional, relatively compact voltage sources. For example, a voltage source such as the voltage sources disclosed in U.S. Pat. Nos. 5,282,122 and 4,241,360, herein entirely incorporated by reference and to which the reader is directed for further details, may be used in certain circumstances where a miniaturized, flexible, device is desired. 
   Each up-converter module provides an incremental voltage up-conversion from an initial input voltage. Such an initial input voltage may be provided from the power source  56  or as an up-converted voltage from another up-converter module. A final up-converted voltage is then available at the X-ray emitter  64 . For example, the first voltage up-converter module  62 ( a ) receives a first input voltage and up-converts this first input voltage to a first output voltage. This first input voltage may be received by the power source  56  of the instrumentation control  54 . Alternatively, because of the current and voltage characteristics of the solid state components making up the up-converter module  62 ( a ), an initial input voltage may not be required. In such a scenario, a photon source  72  is provided by the light source  58  along the fiber optic cable  70 . In this manner, the photons provided by the first fiber optic cable  72  are used to optically generate an output voltage so as to provide an input to the second up-converting module  62   b.    
   The first output voltage (and now a second input voltage) is then applied to the second voltage up-converter module  62 ( b ). This second up-converter module  62 ( b ) up-converts this input voltage to a second output voltage (i.e., a third input voltage). Voltage up-converter  62 ( b ) then provides an up-converted output voltage to a third voltage up-converter module  62 ( c ). As with the first and the second voltage up-converters  62 ( a-b ), the third voltage up-converter module  62 ( c ) up-converters this input voltage and provides an output voltage to the fourth and final power supply module  62 ( d ). In this up-converting manner, the modules may be fabricated so as to produce a known and desired, final up-converted voltage to the X-ray device  64 . This up-converted voltage is then used by the X-ray device  64  to generate the X-rays  66 . 
   An anode of the X-ray emitter receives this voltage from the fourth voltage up-converter and, under the control of the operator control system  57 , emits an X-ray pattern  66  as previously described above. 
   Preferably, both the X-ray source  64  and the up-converting modules are contained within a single enclosure, such as a medical instrument (a catheter). In the arrangement illustrated in  FIG. 3 , four up-converting modules are provided. However, as those of ordinary skill will recognize, other up-converting module arrangements may also be provided. Those of ordinary skill will also recognize, as will be described, various aspects of up-converting module fabrication will tend to effect a number of modules required to eventually produce the necessary and desired final voltage to be provided to the X-ray device  64 . Varying an initial input voltage will also affect the final up-converted output voltage. 
   The up-converting device may be designed to produce a wide array of different voltages. For example, in one arrangement, the first modular section  62 ( a ) has an input potential supplied along input line  71  of approximately 0 to 1000 volts and up-converts this input voltage to approximately 4 kV. This initial input potential could be provided by the power source  56 . Other input potentials could also be provided. For example, in one arrangement, the input potential may be 0 volts. In such an arrangement, the first modular section relies on a light source (photon source via fiber optic cable) to provide an initial voltage up-conversion. 
   In one arrangement, the voltage up-converter modules  62 ( a-d ) are all essentially identical modules. That is, each up-converter module has been fabricated so as to produce essentially the same up-converting characteristics (each module up-converts an input voltage by the same amount: 4-5 kV). Alternative arrangements may also be provided wherein the modules comprise different up-converting characteristics (up-converting rates) to thereby produce different up-converting voltages. For example, a first up-converting module could up-convert an input of 0.1 kV to 3 kV (an up-converting rate of approximately 3 kV) and a second up-converting module could up-convert 3 kV to 9 kV (an up-converting rate of approximately 6 kV). As those of ordinary skill in the art will recognize, other upconverting rates may also be provided. 
   Returning to  FIG. 3 , the second modular section  62   b  receives the output of the up-converting module  62 ( a ) along voltage line  75  and up-converts this input voltage (4-5 kV) to a second voltage that may be provided at voltage line  77 . In one arrangement, this second voltage is 10 kV. This up-converting process is repeated through the remaining modular sections. In this manner, the fourth modular section  62 ( d ) up-converts an input voltage provided along voltage line  79  to an output voltage of 20 kV. For certain angioplasty operations, this is a sufficient voltage. 
   One advantage of the device illustrated in  FIG. 3  is that there is generally only one general location in the entire system where a peak voltage of 20 kV is provided. This point is located at the output of the fourth up-converter modular section  62 ( d ). It is only at the fourth modular section output (at the X-ray unit  64 ), therefore, that the greatest probability of a dielectric breakdown can occur. However, unlike in certain proposed miniaturized high voltage configurations, this peak voltage is not present, nor is it required, along the entire length of the voltage source enclosure. Rather, any peak voltage is provided at only one point: the input of the X-ray unit. Consequently, the overall structural charge-up capacitance of the entire structure is reduced and may be reduced to approximately the size of the last modular section (i.e., the size of modular section  62 ( d )). 
   In one arrangement, the size of the last modular section, and therefore the relevant capacitance, is roughly on the order of about 1 millimeter. In one arrangement, the overall capacitance may be reduced by a factor of 1000 over the proposed system illustrated in  FIG. 1  by reducing the length from 3 feet (˜1000 mm) to 1 mm, which is the X-ray head. Therefore, in the advent of an inadvertent discharge, only 10 Watts would be discharged as compared to 10,000 Watts as mentioned above. 
   The voltage up-converter arrangement illustrated in  FIG. 3  provides a number of advantages. One advantage is that peak voltages are only present at the desired point of X-ray application. That is, the peak voltage is available only near at the X-ray unit. Therefore, the peak voltage need not propagate along the entire length of the catheter. Therefore, the point of highest voltage has the largest probability of dielectric breakdown. Here, because the size of the last modular section is miniaturized, the system&#39;s overall capacitance is also quite small since the length of the capacitance (i.e., the length of the last modular component) is only on the order of 1 mm. Consequently, the overall system concern for flashback is substantially reduced. 
   Another advantage of the arrangement illustrated in  FIG. 3  is its flexible characteristics. As previously discussed, there is a need for a flexible and maneuverable device that allows an up-converted voltage to be applied at certain small locations. Because of the multi-sectioned structure of the arrangement illustrated in  FIG. 3 , the device  50  can be manipulated in various configurations. 
   Several methods may be implemented to fabricate one of the modular sections provided in the system illustrated in FIG.  3 . For example,  FIG. 4  illustrates an arrangement  90  of one of the up-converting modular sections illustrated in FIG.  3 . In this arrangement, the modular section comprises a solid-state device containing a large number (several thousand) laterally fabricated photodiodes. A schematic representation of such a potential photo diode arrangement  90  is illustrated in FIG.  4 . In this schematic representation, the photo diode arrangement  90  has a width of W and a length of L. In one arrangement, this width W is about 1.0 mm and this length L is about 1.3 mm. Such dimensions make this photo diode module an ideal candidate for applications requiring a miniaturized “point of use” power source. 
   As shown in  FIG. 4 , the arrangement  90  includes a serial array of diodes provided along a substrate surface. Preferably, this substrate comprises a sapphire supporting structure. 
   The photodiodes making up the modular section  90  are fabricated in a cascaded, serial fashion. Preferably, and as will be discussed in greater detail below, the photodiodes are laterally disposed in a pattern along a substrate surface and configured in a generally meandering type of configuration as shown in FIG.  4 . The photodiodes begin at a first termination point  93 , wind along the meandering photodiode pattern, and eventually end at a second termination point  95 . 
   The modular section  90  is provided with an input voltage at a modular section voltage input line  92 , up-converts this input voltage, and then provides a modular section voltage output at line  94 . The modular section input, normally a wire, can receive an initial potential voltage (e.g., 0-1000 volts). 
   Line  94  provides an output voltage. Where the input wire supplies an input voltage, the modular section up-converts the initial potential voltage to a second potential voltage which can then be supplied as an input to another modular section. Alternatively, where the modular structure  90  is the last modular section in a cascaded plurality of sections (such as the forth modular section  62 ( d ) illustrated in FIG.  3 ), section  90  provides an up-converted peak voltage to a device, such as the X-ray unit  64  illustrated in FIG.  3 . 
   The modular section  90  is optically coupled to a light source via an optical fiber  96 . Fiber  96  provides a source of light (photons) so as to energize the plurality of photo diodes  100 . For example, in one arrangement, the modular section  90  receives a source of photons  97  over optical fiber  96 , wherein, the optical fiber  96  is optically coupled to a light source, such as the light source  58  illustrated in FIG.  3 . In one arrangement, the optical fiber  96  has a diameter of 120 μm (0.12 mm). In one arrangement, the fiber optic cable  96  is bundled with a plurality of other optical fibers. These various optical fibers act as a separate photon source to each modular section. 
   The modular section can be fabricated to have a length designated L and a width designated W such that the modular section is small enough to be contained in a miniature instrument, such as the instrument illustrated in FIG.  3 . More preferably, in one arrangement, the designated length L is 1.3 mm and the designated width W is 1 mm. However, as those of ordinary skill in the art will recognize, other structures, configurations, and/or dimensions may also be utilized. 
   The modular structure  90  may be encapsulated within an encapsulation media  99 . Encapsulation media  99  is shown in  FIG. 4  as surrounding or “encapsulating” the modular device substrate. Wires  92 ,  94  and the fiber optic cable  96  protrude outside the encapsulated area. In such an arrangement, the encapsulation media  99  provides a degree of optical and electrical isolation between the optical sensitivities of the plurality of photodiodes  100  and an environment surrounding the encapsulated media. The media could comprise certain plastics, pyrelene, Teflon, polyimide, certain forms of polydimethylsiloxane (PDMS), or other types. The encapsulation medium  99  also provides a degree of stability (or support) for the wire  92 , the outgoing wire  94 , and the fiber optic cable  96 . 
   Utilizing the arrangement illustrated in  FIG. 4 , a large number of photo-diodes may be fabricated onto a small substrate footprint. Generally, the greater the number of photodiodes per module, the greater the module&#39;s up-converting rate. This can be illustrated by equating the number of diodes that may be fabricated onto a 1×1.3 mm 2  chip, such as the modular section illustrated in FIG.  4 . The unit cell (area per diode) is 15×10 μm 2  so this equals 150 μm 2 . The total area of the up-converter module is 1×1.3 mm=1000×1300=1.3×10 6  μm 2 . If N is defined as the number of diodes per module, one can see that N=(1.3×10 6 )/150 which equals approximately 8,700 diodes. Therefore, if each photodiode generates a photo voltage of approximately 0.5 volts, a module comprising 8,700 photo-diodes can generate 4350 volts (4.35 kV). Modular sections having other up-converting rates may also be fabricated in a similar manner. In addition to silicon as described below, other photo materials may also be used, including Gallium Arsenide. 
   The photodiodes provided in the modular section  90  may be fabricated utilizing various methods. One such method involves the fabricating process  112  illustrated in FIG.  5 . Process  112  is particularly useful in fabricating an array of laterally disposed photo-diodes. Process  112  will be described in reference to FIG.  5  and the various steps illustrated in FIGS.  6 ( a-i ). 
   First, at step  114  and as illustrated in FIG.  6 ( a ), a silicon-on-insulator substrate  142  may be quartz (SiO 2 ) or alternatively, sapphire (Al 2 O 3 ), is first provided. This substrate contains commercially available polycrystalline silicon, laser crystallized polycrystalline silicon, or single crystal silicon. This is referred to as Silicon-On-Insulator (S-O-I). 
   Ion implantation (phosphorus or arsenic) is performed to render the undoped silicon of the SOI wafer slightly n −  type at a desired concentration, preferably at a concentration of approximately 1×10 15  atoms/cm 3  to 8×10 15  atoms/cm 3 . 
   Next, referring now to FIG.  6 ( a ), a photo-diode pattern  144  is formed by depositing a positive photoresist on substrate  142 . This photo-diode patterning occurs at step  118  in FIG.  5 . 
   Next, a photoresist is provided. Such a positive photoresist may be Shipley 1818 that is deposited onto the substrate  142  by spin coating. Other coating methods could also be used. After baking at about 100° C. for several minutes, the photoresist is then exposed via a mask using an ultraviolet light source. The photodiode pattern  144  is developed and the silicon is etched using a plasma etcher containing a CF 4 /O 2  mixture. Other etching gases can be used as well as chemical etchants. 
   The pattern  144  is preferably of a meander-type pattern. As will be explained in further detail below, such a meandering photo-diode type pattern results in an array structure that provides a high density of serially connected laterally constructed diodes (photo-diodes/mm 2 ). Those of skill in the art, however, will note that other type of photo-diode pattern could also be used. For example, certain patterns could be chosen that maximize the distance between the input voltage and the up-converted voltage. 
   As can be seen from FIG.  6 ( a ), the meander-type pattern  144  comprises a number of columns  145 ( a-e ) and a number of rows  141 ( a-e ). These various columns  145 ( a-e ) are attached via a number of rows  141 ( a-e ). For example, column  145 ( c ) is connected to column  145 ( d ) via row  141 ( c ). In one arrangement, each column has a width w 1  of approximately 5 μm and each row connecting adjacent columns has a width w 2  of approximately 5 μm. 
   The meandering type pattern  144  begins at a first terminal point  156  and extends across a surface  142  of the substrate  140  to a second terminal point  157 . The first terminal point  156  has a larger width than the rows and will preferably provide a contact point for an electrical connection, such as for the wire  92  illustrated in FIG.  4 . Similarly, the second terminal point  157  has a width large enough so as to provide a contact point for another electrical connection, such as for the wire  94  illustrated in FIG.  4 . 
   As a next step in the fabrication process, a p +  type implant takes place. (step  120 , FIG.  5 ). FIG.  6 ( b ) illustrates p +  type implantation along a portion of the photo-diode pattern  144  illustrated in FIG.  6 ( a ). Prior to this procedure, the photoresist on top of the n −  silicon  144  of FIG.  6 ( a ) has been removed via a wet or a dry stripping step. 
   A photo-resist step is now repeated to fabricate a multitude of p +  regions along the meandering pattern  144 . Boron could be used to fabricate these p +  regions. A photo resist is then spun on, baked, exposed, and developed and boron is ion implanted to a concentration of about 1×10 18  ions/cm 3  to about 5×10 18  ions/cm 3  in the regions where the photoresist has been developed (see  FIG. 6   b ). 
   FIG.  6 ( b ) illustrates a fabricated region portion  150  of the n −  implanted meander  144 . As illustrated in FIG.  6 ( b ), the fabricated portion  150  comprises a first p +  implant zone  156  and a second p +  implant zone  158 . Adjacent this first p +  implant region  156  is a first photoresist  154  protecting an underlying n −  implanted region during p +  implant. The second p +  implant region  158  is provided adjacent the first and the second photo resists  154 ,  152 , respectively. The remainder of the entire photo-diode pattern  144  of FIG.  6 ( a ) extending from the first termination point  156  to the second termination point  157  is fabricated in a similar manner. 
   Prior to the next implantation step, the photoresist regions are removed. (step  122  in  FIG. 5 ) and an n +  implantation process occurs at step  126  of FIG.  5 . This n +  implantation step  126  is illustrated in FIG.  6 ( c ). As shown, fabricated substrate portion  160  includes a first, a second, and a third photoresist area  162 ,  164 , and  166 , respectively. These photoresist areas act to protect the underlying previously implanted regions. The n +  implants are represented by areas  170  and  168 . In one arrangement, the n +  implant comprises either phosphorous or arsenic and is implanted to a concentration of approximately 1×10 18  to 5×10 18  ions/cm 3 . 
   In a next step, the photoresist regions  162 ,  164 , and  166  are removed. Removing the photoresists  162 ,  164 , and  166  results in a top view of a portion of a fabricated device  180  is illustrated in FIG.  6 ( d ). As shown in FIG.  6 ( d ), the fabricated device  180  now comprises a p +  region defining a first termination point  182 . Adjacent this p +  region termination point  182  is an n− type region  183 A, an n+ type region  183 B, and then another p+ type region  183 C. This p+ to n −  to n +  pattern is repeated throughout the photodiode pattern, extending from the first termination point  156  to the second termination point  157  of FIG.  6 ( a ). 
   At step  128  (FIG.  5 ), the silicon and substrate are cleaned. After the cleaning step  128 , a protective oxide layer, preferably SiO 2 , is grown over the various doped silicon regions. This occurs at step  130 . In one arrangement, an oxide layer of about 1000 Angstroms is grown at about 950-1000 degrees Celsius. During the oxide layer growing process, the ion implanted regions are being activated, i.e., the ion implanted regions are rendered electrically conductive. 
   FIG.  6 ( e ) illustrates a cut away view along view B—B′ of the substrate portion  180  illustrated in FIG.  6 ( d ). This cut-away view illustrates the substrate  194  after an oxide layer  192  has been grown over the surface  193  of the device  190 . As shown, the lateral array of the various doped regions comprising the photodiodes  197  are disposed along the top surface  193  of the substrate layer  194 . 
   A next process step includes photo-masking the substrate to form a contact opening. Preferably, at least two contact openings per photodiode module are formed. For example, as illustrated in FIG.  6 ( f ), a photo-resist  202  is provided over a portion of the substrate, excluding the developed areas over region  182  of  FIG. 6   d  and partial regions where p+ and n+ regions meet. The SiO 2  layer, in the unprotected regions, is etched using buffered HF (hydrofluoric acid). After photoresist stripping, the device looks like what is illustrated in  FIG. 6   g . The SiO 2  is removed from surface portions  211 ,  213  overlaying the laterally disposed p+ region  211  and from a region overlapping adjacent n+ and p+ regions  213 . 
   In the next process step  131  (FIG.  5 ), contact material is deposited along the surface of the fabricated substrate. This deposited contact material is photo-shaped to form a contact region for the voltage input and output. This deposited contact material is also photo-shaped to shunt the p+ and the n+ regions. Preferably, the contact material is aluminum, however, Cr/Au or other similar materials. Good contacts can be formed to the silicon regions and wire bonds may be formed on the contact material. FIG.  6 ( h ) shows a top view of the photodiode device with two metal regions  242 ,  244  as described above. Wire  264  may then be connected to contact material in FIG.  6 ( i ) by wire bonding, such as to apply an input voltage for up-converting. 
   FIG.  6 ( i ) illustrates the cross section of the device shown in FIG.  6 ( h ) with the inclusion of the wire to  264  to contact material  256 . 
   As a next processing step, the device illustrated in FIG.  6 ( i ) may be encapsulated in an encapsulation media. Encapsulation provides a number of advantageous features. For example, encapsulation protects against moisture, provides the fabricated up-converting module with an enhanced level of rigidity, and also prepares for inclusion into a medical device. 
   The process illustrated in FIGS.  5  and  6 ( a-i ) results in a plurality of photodiodes laterally fabricated along a substrate surface.  FIG. 7  illustrates a current versus voltage graph  270  that demonstrates how the fabricated photo-diode device utilizes an optical source to provide voltage up-converting. As shown in the graph  270 , the line V dark    278  represents the ordinary operating characteristic of a photodiode absent any illumination. In this condition, the voltage versus current characteristics demonstrate that there is no voltage output at zero current. 
   However, once the photodiode is illuminated, the current versus voltage graph shifts. This shift is graphically represented as V light ,  277 . V light    277  has now shifted along the y-axis so that now, the graph intercept with the x-axis has shifted where this shift is defined as V OC  or the photodiode&#39;s photo-voltage when no current is drawn (an open circuit condition). Ordinarily, such a diode photo-voltage may be on the order of approximately 0.5 volts. When current is drawn, the open circuit voltage is somewhat lowered. 
     FIG. 8  illustrates a perspective view of a fabricated miniaturized up-converter module  300  prior to encapsulation. The miniaturized up-converter device  300  comprises the doped regions arranged in a meandering pattern  309  and provided along a substrate surface  302 . A first wire  306  provides an input voltage V 1  at a first termination point. A second wire  304  is used to supply a remote device with an up-converted output voltage V 2  at the second termination point. This upconverted output voltage V 2  may then be supplied to another miniaturized up-converter module, like device  300 , or may be used to provide an up-converted voltage to a device such as an X-ray device. 
     FIG. 9  illustrates a photodiode module  320  connected to a fiber. The encapsulated module  320  provides a photodiode device  346  encapsulated within a structure  348 . The photodiode device  346  is encapsulated within a structure  348 . The photodiode device  346  includes an input wire  322  and an output wire  324 , both wires extend beyond the encapsulation structure  348 . At one end of the enclosure, a fiber optic light source  326  is provided for providing a source of photons  328 . These photons are incident along the plurality of photodiodes residing along a bottom surface  344  of the up-converting module  346 . As the photons propagate along the length of the module  346 , from the fiber optic cable  326  to a back end  342 . Some of the photons reflect off the module bottom surface  344  and off of a bottom enclosure portion  341 . The fiber optic cable  326  is fixedly attached to the module via glue  342  or some other adhesive. 
   A reflective surface  340  may be provided along the top surface of this portion so as to increase an overall reflectivity along the bottom portion  341 . A structure back end  342  may also be provided with a reflecting medium  343 . The reflective medium may be chosen to have an index of reflection so as to enable the photons reflecting off of this back surface to be totally internally reflected. 
   In one arrangement, the up-converting module, such as the module illustrated in  FIG. 9 , may be encapsulated. For example, two types of module encapsulation arrangements are illustrated  FIGS. 10   a  and  b .  FIG. 10   a  illustrates one encapsulation arrangement wherein each separate module is independently encapsulated. In  FIG. 10   a , a first encapsulated module  361  is provided with an input wire  364  and an output wire  366 . The output wire  366  is electrically coupled to an input wire of a second encapsulated module  363 . The second encapsulated model  363  also includes an output wire  362  that may be electrically coupled to another encapsulated module. Alternatively, output wire  362  may be electrically coupled to a device requiring a peak up-converted voltage, such as an X-ray device. 
     FIG. 10   b  illustrates an alternative encapsulation arrangement.  FIG. 10   b  illustrates an encapsulation arrangement wherein two up-converting modules are encapsulated within a single encapsulation structure  378 . In  FIG. 10   b , a first module  371  has an input wire  372  and an output wire  376 . The output wire  376  is coupled to an input wire of a second module  373 . The second model  373  includes an output wire  374  that may be coupled to another encapsulated module or other device requiring an up-converted voltage, such as an X-ray device. 
   Exemplary embodiments of the present invention have been described. Those skilled in the art will understand, however, that changes and modifications may be made to these embodiments without departing from the true scope and spirit of the present invention, which is defined by the claims.