Patent Publication Number: US-9849296-B2

Title: Directly integrated feedthrough to implantable medical device housing

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     This application is a divisional of U.S. patent application Ser. No. 14/104,653, filed on Dec. 12, 2013, entitled “DIRECT INTEGRATION OF FEEDTHROUGH TO IMPLANTABLE MEDICAL DEVICE HOUSING BY SINTERING,” now U.S. Pat. No. 9,610,452, to be issued Apr. 4, 2017, which is incorporated herein by reference. 
     This Patent Application is related to Ser. No. 14/104,636, filed on Dec. 12, 2013, entitled “DIRECT INTEGRATION OF FEEDTHROUGH TO IMPLANTABLE MEDICAL DEVICE HOUSING USING A GOLD ALLOY,” now U.S. Pat. No. 9,610,451 to be issued Apr. 4, 2017, and U.S. Pat. No. 9,504,841, issued Nov. 29, 2016, entitled “DIRECT INTEGRATION OF FEEDTHROUGH TO IMPLANTABLE MEDICAL DEVICE HOUSING WITH ULTRASONIC WELDING,” all of which are incorporated herein by reference. 
    
    
     BACKGROUND 
     Implantable medical devices, such as cardiac pacemakers, cardiac defibrillators, and neurostimulators, receive and/or deliver electrical signals to/from portions of the body via sensing and/or stimulating leads. Implantable medical devices typically include a metal housing (typically titanium) having a hermetically sealed interior space which isolates the internal circuitry, connections, power sources, and other device components from body fluids. A feedthrough device (often referred to simply as a feedthrough) establishes electrical connections between the hermetically sealed interior space and the exterior bodily fluid side of the device. 
     Feedthroughs typically include an insulator (typically ceramic) and electrical conductors or feedthrough pins which extend through the insulator to provide electrical pathways between the exterior and the hermetically sealed interior. A frame-like metal ferrule is disposed about a perimeter surface of the insulator, with the ferrule and insulator typically being joined to one another via a brazing or soldering process. The ferrule is configured to fit into a corresponding opening in the metal housing, with the ferrule being mechanically and hermetically attached to the housing, typically via laser welding. The insulator electrically insulates the feedthrough pins from one another and from the metal ferrule/housing. 
     The ferrule is typically joined to insulator via a welding or brazing process. However, the high temperatures employed by such processes heats the titanium of the housing about the perimeter of the opening to levels that cause a structural change in the titanium, commonly referred to as “grain growth”. This structural change can distort the dimensions of the opening and cause the titanium about the perimeter of the opening to become less rigid, each of which can result in a weaker joint between the ferrule and the housing. 
     Additionally, machining the ferrule (typically from pure titanium) to provide a high tolerance gap between the ferrule and the insulator (about 10-50 μm) which is necessary to achieve a quality braze joint is demanding and costly. Furthermore, if the gap is not maintained during the brazing process, or if the brazing process itself is not properly performed, a weak joint may be formed that can lead to premature failure of the implantable device. 
     For these and other reasons there is a need for the embodiments of the present disclosure. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
       The accompanying drawings are included to provide a further understanding of embodiments and are incorporated in and constitute a part of this specification. The drawings illustrate embodiments and together with the description serve to explain principles of embodiments. Other embodiments and many of the intended advantages of embodiments will be readily appreciated as they become better understood by reference to the following detailed description. The elements of the drawings are not necessarily to scale relative to each other. Like reference numerals designate corresponding similar parts. 
         FIG. 1  generally illustrates an example of an implantable medical device according to one embodiment. 
         FIG. 2  illustrates a feedthrough device in an implantable in accordance with the prior art. 
         FIG. 3  illustrates a cross-sectional view of a feedthrough in an implantable medical device in accordance with one embodiment. 
         FIG. 4  is cross-sectional view illustrating a feedthrough in an implantable medical device including according to one embodiment. 
         FIG. 5  is cross-sectional view illustrating a feedthrough in an implantable medical device including according to one embodiment. 
         FIG. 6  is a block and schematic diagram illustrating a method of attaching a feedthrough to a housing using sintering process according to one embodiment. 
         FIG. 7  is a feedthrough to a housing using sintering process according to one embodiment. 
         FIG. 8  is a feedthrough to a housing using sintering process according to one embodiment. 
         FIG. 9  is a flow diagram illustrating a method of attaching a feedthrough to a housing using sintering according to one embodiment. 
     
    
    
     DETAILED DESCRIPTION 
     In the following Detailed Description, reference is made to the accompanying drawings, which form a part hereof, and in which is shown by way of illustration specific embodiments in which the invention may be practiced. In this regard, directional terminology, such as “top,” “bottom,” “front,” “back,” “leading,” “trailing,” etc., is used with reference to the orientation of the Figure(s) being described. Because components of embodiments can be positioned in a number of different orientations, the directional terminology is used for purposes of illustration and is in no way limiting. It is to be understood that other embodiments may be utilized and structural or logical changes may be made without departing from the scope of the present invention. The following detailed description, therefore, is not to be taken in a limiting sense, and the scope of the present invention is defined by the appended claims. 
     It is to be understood that the features of the various exemplary embodiments described herein may be combined with each other, unless specifically noted otherwise. 
     In accordance with one embodiment of the present disclosure, a method of attaching a feedthrough to a titanium housing of an implantable medical device is provided. The method includes applying a sinter paste onto a surface of the housing about a perimeter of an opening through the housing, the sinter paste including a biocompatible bonding material. An insulator of the feedthrough is placed onto the sinter paste so as to cover the opening, and the sinter paste is heated to a temperature less than a beta-transus temperature the titanium of the housing and to a temperature less than a melting point of the biocompatible bonding material for a desired duration to form, from the sinter paste, a sinter joint which bonds the feedthrough to the housing and hermetically seals the opening. According to one embodiment, portions of surfaces of the insulator contacting the sinter paste and resulting sinter joint are metallized. 
     Embodiments described herein for sintering the insulator of a feedthrough device directly to the device housing provides advantages over known processes of attaching a feedthrough device to device housing. First, attaching the feedthrough directly to the housing using a sinter joint eliminates the need for a ferrule (such as ferrule  56  of  FIG. 2 ). By directly integrating the feedthrough to the housing via a sinter joint, as opposed to conventional techniques which integrate the feedthrough to the housing using a ferrule, the shortcomings associated with such a ferrule (e.g. brazed/welded joint, machining requirements, costs) are eliminated. Additionally, when combined with the use of cermet for conductive elements of the feedthrough, the present disclosure provides a feedthrough which is completely devoid of welds and/or brazing. 
     Additionally, by using a sintering process as described herein at temperatures below the β-transus temperature of titanium of the device housing, grain growth within the titanium material of the housing is greatly reduced relative to conventional techniques which employ high-temperature brazing or welding processes to attach feedthrough devices to housings via a ferrule, particularly in the region of the housing about a perimeter of an opening in which the feedthrough is disposed. In one embodiment, the temperature does not exceed 750° C. In one embodiment, the temperature does not exceed 350° C. Reducing the grain growth of the titanium of housing reduces dimensional distortions of housing as compared to conventional techniques, at least to levels within design tolerances, thereby providing stronger and more consistent hermetic seals between the insulator and the housing. 
     According to one embodiment, the sinter paste is formed by mixing the biocompatible bonding material in a powdered form with a binder material. In one embodiment, the powdered biocompatible bonding material has particles with a maximum dimension of less than 20 μm. In one embodiment, the particles are spherical in shape with a diameter less than 20 μm. In one embodiment, the biocompatible bonding material comprises gold. In one embodiment, the biocompatible bonding material comprises one of gold, platinum, palladium, and any alloy combination thereof. 
     According to one embodiment, the method includes applying a force to the feedthrough during the heating to push the insulator toward the housing so as to compress the sinter paste as the sinter paste loses volume due to binders within the sinter paste burning off during the heating, thereby providing a stronger joint and hermetic seal. In one embodiment, a counter force is provided to support the housing to prevent deflection of the housing. In one embodiment, the heating is carried out in an oven having a non-oxygen and non-vacuum environment to enable the melted sinter material to flow and to prevent oxidation of the titanium material of the housing, thereby enabling a stronger joint and improved hermetic seal. 
     Another aspect provides an implantable medical device including a housing having an opening with an opening width, and a feedthrough including an insulator having a bottom surface and side surfaces and having an insulator width between opposing side surfaces that is greater than the opening width. A sinter joint between at least one of the bottom surface, top surface, and side surfaces of the insulator and the housing hermetically seals the insulator to the housing. 
     In one embodiment, a width of the sinter joint between the insulator and the housing along a perimeter of the housing opening is at least one quarter a width of the insulator at its widest point. In one embodiment, the housing includes a flange which forms a recess about the opening, the opening being disposed at a bottom of the recess, and the insulator being positioned at least partially within the recess and over the opening. In one embodiment, a plane normal to the opening passes through the housing, the sinter joint, and the insulator. In one embodiment, the sinter joint has a thickness in a direction perpendicular to the housing in a range from 25 to 200 μm. 
     In one embodiment, the sinter joint comprises gold. In one embodiment, the sinter joint comprises one of gold, palladium, iridium, and alloy combinations thereof. In one embodiment, the sinter joint has a density of not more than 99 percent of the biocompatible bonding material. In one embodiment, the sinter joint has a density in a range from 90 to 99 percent of the biocompatible bonding material. In one embodiment, the housing comprises titanium, and wherein the titanium has an average grain size of not more than 100 μm. 
     A further aspect of the present disclosure provides a method of attaching a feedthrough device to hermetically seal an opening in a titanium housing of an implantable medical device. A sinter paste is applied about a perimeter of the opening, the sinter paste including a biocompatible bonding material, and a feedthrough is positioned on the sinter paste to cover the opening, the feedthrough having a width greater than a width of the opening. The sinter paste is heated to a temperature less than a β-transus temperature of the titanium of the housing and less than a melting point of the biocompatible bonding material for a duration which to limit an average grain size of the titanium to not greater than 100 μm and to form a sinter joint from the sinter paste that bonds the feedthrough to the housing and hermetically seals the opening. As described above, by reducing the grain growth of the titanium of housing relative to conventional processes, which result in grain sizes well in excess of 100 μm (for example, greater than 300 μm), dimensional distortions of housing are reduced as compared to conventional techniques, at least to levels within design tolerances, thereby providing stronger and more consistent hermetic seals between the insulator and the housing. 
     In one embodiment, the method includes forming the sinter paste by mixing the biocompatible bonding material in a powder form with a binding material. In one embodiment, the method includes applying a force to the feedthrough during the heating to push the insulator toward the housing so as to compress the sinter paste as the sinter paste loses volume due to binders within the sinter paste burning off during the heating, and providing a counter force to support the housing to prevent deflection of the housing. 
       FIG. 1  is a block and schematic diagram generally illustrating one embodiment of an implantable medical device  30 , such as a cardiac pacemaker for example. Implantable medical device  30  includes a hermetically sealed metal case our housing  32 , typically formed of titanium, which defines a hermetically sealed interior space  34  in which device electronics  36  are disposed and protected from fluids of the body fluid side  38  external to housing  32 . A header  40  attaches to housing  32  and includes a connector block  42  which typically includes one or more sockets for connecting to one or more sensing and/or stimulating leads  44  that extend between implantable medical device  30  and desired regions of the body, such as the human heart and brain, for example. A feedthrough device  50  establishes electrical pathways or connections through housing  32  that maintain the integrity of hermetically sealed interior space  34  and provide electrical connection of leads  44  to internal device electronics  36 . 
       FIG. 2  is a cross-sectional view illustrating portions of an implantable medical device, such as medical device  30  of  FIG. 1 , including metal housing  32  having an opening  46  in which a conventional feedthrough device  50  is positioned. Feedthrough device  50  includes an insulator  52 , feedthrough pins or conducting elements  54 , and a ferrule  56 . A ferrule  56 , comprising a frame-like metal structure, holds insulator  52  and which is configured to fit into opening  46  for attachment to housing  32 . Ferrule  56  is a bio-compatible material, typically titanium, which is mechanically and hermetically attached to housing  32  by laser welds  58 , or similar techniques. Ferrule  56 , as illustrated in  FIG. 2 , sometimes includes a flange  60  to further aid in securing ferrule  56  to housing  32 . 
     Conducting elements  54  extend through openings or vias  62  in insulator  52  and are formed of an electrically conductive material so as to provide electrically conductive pathways from the external body fluid side  38  of housing  32  to hermetically sealed interior space  34 . Insulator  52  is formed of a non-electrically conductive material, such as a ceramic material, aluminum oxide (Al 2 O 3 ) for example, and electrically isolates conducting elements  54  from one another and from ferrule  56  and housing  32 . 
     When attaching insulator  52  and ferrule  56  to one another, a perimeter surface of insulator  52  is typically metalized (through a sputter coating process, for example) to provide a thin metal coating  64  thereon. Ferrule  56  is then joined to insulator  52  via metal coating  64  using a braze  66 , such as of gold, for example, to form a biocompatible and hermetic seal. Similarly, interior surface of vias  62  are provided with a metal coating  68  and a braze  70  (e.g. gold) is used to couple conducting elements  54  to insulator  52  and form a biocompatible and hermetic seal. 
     In order to achieve a quality braze, and thereby a quality hermetic seal, a proper gap must be maintained between ferrule  56  and insulator  52  during the brazing process (typically about 10-50 μm) so that the brazing material (e.g. gold) is properly drawn into the gap by capillary action to create a strong and reliable braze  66 . Forming ferrule  56 , typically via machining processes, to meet the tight tolerances required to provide the proper gap with insulator  52  as well as to the dimensions of opening  46  in housing  42  is time consuming and costly. Also, during the brazing process, intermetallics are formed between the brazing material (e.g. gold) and the material (e.g. titanium) of ferrule  56 , with the intermetallics being brittle as compared to the brazing material. If the gap between ferrule  56  and insulator  52  is too small, the amount of intermetallics may be large relative to the amount of pure brazing material (e.g. gold) resulting in a brittle braze  66  that may crack and comprise the hermitic seal. 
     Additionally, heat from the brazing (or welding) of ferrule  56  to housing  32  can cause structural changes in the titanium of housing  32  about opening  46  (and to ferrule  56 ) due to “grain growth” in the titanium. Such “grain growth” can cause undesirable dimensional changes in opening  46  and can cause the titanium about the perimeter of opening  46  to become less rigid (i.e. more flexible), which such changes leading to a weakened or defective joint. 
     All polycrystalline materials, including titanium, are made of closely packed atoms, with “regions of regularity” within these closely packed atoms (i.e. where the atoms have a regular structure, such as 8-co-ordination and 12-co-ordination, for example) being referred to as “crystal grains”. Metal consists of a vast number of these crystal grains. The boundaries of these crystals (i.e. “grain boundaries”) are locations at which atoms have become misaligned (i.e. the regular structure is discontinuous). Metals having smaller grains and, thus, more grain boundaries, are harder than metals having larger grains, which have fewer grain boundaries and, as a result, are softer and more flexible. 
     Heating of a metal, such as titanium, causes the atoms to move into a more regular arrangement, thereby decreasing the overall number of crystal grains but increasing the grain size of the remaining grains (i.e. the number of grains per unit volume decreases). The process by which the average grain size increases, so-called “grain growth”, rearranges the crystalline structure of the metal and can cause dimensional changes (i.e. dimensional deformation) of the metal and cause the metal to become more flexible. 
     Titanium has an α-phase, which has a close-packed hexagonal crystal structure, and a β-phase, which has centered-cubic crystal structure and that is more open and prone to grain growth than the hexagonal structure. Titanium transitions from α-phase to β-phase, the so-called β-transus, when heated to or above a certain temperature, referred to as the β-transus temperature. The β-transus temperature is affected by impurities in the titanium (e.g. iron, carbon, hydrogen), but typically occurs at about 880° C. in commercially-pure titanium. Commercially pure titanium, as opposed to titanium alloys having additive such as aluminum (Al), typically has a microstructure of primarily α-phase grains having an average grain size in the range of 10-40 μm. 
     The grain growth of a metal, including titanium, is a function of the time and temperature for which a metal is heated. For example, while the average grain size of commercially-pure titanium increases when heated to temperatures below the β-transus temperature, such grain growth accelerates rapidly when the the titanium is heated to a temperature at or above the β-transus temperature and the titanium transitions from α-phase to β-phase. For instance, the average grain size of commercially-pure titanium has been shown to increase in from about 10-40 μm to about 70 μm when heated at 700° C. for 120 minutes, to about 100 μm when heated at 750° C. for 120 minutes, and to about 180 μm when heated at 800° C. for 120 minutes. However, the average grain size of commercially-pure titanium has been shown to increase in from about 10-40 μm to about 350 μm when heated at 1000° C. for 120 minutes, and to about 425 μm when heated at 1100° C. for 120 minutes. 
     With reference to conventional feedthrough  50  of  FIG. 2 , attaching ferrule  56  to housing  32  by laser welding or brazing (e.g. gold braze) heats housing  32  to a temperature well above the β-transus temperature of titanium, resulting in rapid grain growth in the titanium of housing  32 . For example, the average grain size may increase by 300 μm or more. Such grain growth causes dimensional distortions in housing  32  that can cause opening  46  to be outside of specified tolerances and causes the titanium about the perimeter of opening  46  to become less rigid, each of which can result in a poor or defective seal being formed between housing  32  and feedthrough  50 . 
       FIG. 3  is a schematic diagram illustrating portions of an implantable medical device  130  including a feedthrough  150  according to one embodiment of the present disclosure. Feedthrough  150  includes an insulator  152  and conducting elements  154  extending therethrough. As will be described in greater detail below, feedthrough  150  is attached directly to housing  132  via insulator  152  using a sinter joint  180  that is formed at low-temperatures, at least at temperatures below the β-transus temperature of the titanium of housing  132 . 
     By attaching feedthrough  150  directly to housing  132  via insulator  152 , the need for a ferrule (such as ferrule  56  of  FIG. 2 ) is eliminated, thereby eliminating the cost of manufacturing such a ferrule as well as the difficulties and shortcomings associated with attaching such a ferrule to the insulator (such as insulator  52  of  FIG. 2 ). Additionally, by attaching feedthrough  150  to housing  132  using sintering techniques at reduced temperatures relative to conventional welding or brazing techniques, dimensional distortions of housing  132  due to the high temperatures and grain growth of titanium are substantially reduced, at least to levels that maintain dimensions of housing  32  within specified tolerances, and the titanium remains in a more rigid state. 
     While  FIG. 3  a cross-sectional view illustrating portions housing  132 , particularly the location where feedthrough  150  attaches to housing  132  to seal opening  146 , implantable medical device  130  may include additional features similar to those described with respect to medical device  30  of  FIG. 1 . According to one embodiment, housing  132  is formed of titanium and defines a sealed interior space  134  in which device electronics are disposed and protected from fluids of body fluid side  138  external to housing  132 . According to one embodiment, a header, similar to header  40  of  FIG. 1 , for example, maybe also provided which attaches to housing  132  and includes a connector block having one or more sockets for connecting to one or more sensing and/or stimulating leads. 
     Similar to that described above with regard to  FIG. 3 , feedthrough  150  establishes electrical connections or pathways from body fluid side  138  to the interior space  134  of housing  132  while maintaining the integrity of hermetically sealed interior space  134  via conducting elements  154  which pass through insulator  152 . According to one embodiment, insulator  152  is a glass or ceramic material, such as aluminum oxide (Al 2 O 3 ). According to one embodiment, conducting elements  154  are formed of a cermet. 
     In the context of one embodiment, the terms, “cermet” or “cermet-containing,” refers composite materials made of ceramic materials in a metallic matrix (binding agent). These are characterized by their particularly high hardness and wear resistance. The “cermets” and/or “cermet-containing” substances are cutting materials that are related to hard metals, but contain no tungsten carbide hard metal and are produced by powder metallurgical means. A sintering process for cermets and/or cermet-containing elements proceeds is the same as that for homogeneous powders, except that the metal is compacted more strongly at the same pressuring force as compared to the ceramic material. The cermet-containing bearing element has a higher thermal shock and oxidation resistance than sintered hard metals. In most cases, the ceramic components of the cermet are aluminum oxide (Al 2 O 3 ) and zirconium dioxide (ZrO 2 ), whereas niobium, molybdenum, titanium, cobalt, zirconium, chromium and platinum are conceivable as metallic components. 
     According to one embodiment, such as illustrated by  FIG. 3 , the ceramic (e.g. Al 2 O 3 ) of insulator  152  and the cermet of conducting elements  154  are formed in a first process such that an interface between insulator  152  and conducting elements  154  are hermetically sealed without the use of a braze or solder. According to one example of such an embodiment, the ceramic of insulator  152  is a multi-layer ceramic sheet into which a plurality of vias is introduced. The cermet of conducting elements  154  is then introduced into the vias. In one embodiment, both materials are introduced in a green state, and the combination is fired together. According to such an embodiment, the joining of insulator  152  with conducting elements  154  forms a hermetic seal without the use of braze or solder. 
     According to one embodiment, sinter joint  180  is formed of a biocompatible material, such as gold or a gold alloy, for example, which is applied as a sinter paste prior to the carrying out of a sintering process to form sinter joint  180  (with such process being described in greater detail below). According to one embodiment, the surfaces of insulator  152  at which sinter joint  180  is to be formed are provided with a metallized layer  164  using a suitable process, such as sputter coating or electroplating process, for example. According to one embodiment, metallized layer  164  comprises a biocompatible metal such as niobium, platinum, palladium, titanium, and gold, for example. 
     According to one embodiment, feedthrough  150  has a width W F  at a widest point between opposing surfaces  172   a ,  172   b , which is wider than a width W O  of opening  146  in housing  132 . Insulator  152  further includes an upper surface  174  and a lower surface  176 . It is noted that feedthrough  150  is illustrated in vertical cross-section in  FIG. 3 , but in horizontal cross-section (i.e. between upper and lower surfaces  174  and  176 ) feedthrough  150  can be of a variety of shapes, such a circular, oval, and rectangular, for example. According to one embodiment, sinter joint  180  has a width W J  and a thickness T J . According to one embodiment, housing  132 , feedthrough  150 , and sinter joint  180  are disposed relative to another such that a line  178  drawn through at least a portion of feedthrough  150 , wherein line  178  is orthogonal to upper surface  174  of feedthrough  150 , passes through housing  132 , sinter joint  180 , and feedthrough  150 . 
     According to one embodiment, sinter joint  150  is formed from a biocompatible material. According to one embodiment, sinter joint  150  is formed of one of gold, platinum, palladium, and any alloy combination thereof. According to one embodiment, the thickness T J  of sinter joint  180  is in a range from 20 to 200 μm. 
       FIG. 4  is schematic diagram illustrating implantable medical device  150  according to one embodiment of the present disclosure. According to the embodiment of  FIG. 4 , titanium housing  132  includes a flange  182  that forms a recess  190  in housing  132  about opening  146 , with opening  146  disposed at the bottom of recess  190 . According to one embodiment, as illustrated, flange  182  includes a downwardly angled portion  184  that forms a sidewall of recess  190  and transitions to a horizontal portion  186  that forms a bottom of recess  190  about a perimeter of opening  146 . According to one embodiment, as illustrated, sinter joint  180  is formed between insulator  152  and horizontal portion  186  of flange  182  of housing  132 . 
       FIG. 5  is schematic diagram illustrating implantable medical device  150  according to one embodiment of the present disclosure. According to the embodiment of  FIG. 4 , titanium housing  132  includes flange  182  which forms recess  190  in housing  132 . However, unlike the embodiment of  FIG. 4 , flange  182  includes only a downwardly angled portion  184  and has no horizontal portion such that recess  190  is wider at the top than at the bottom and that an open bottom of recess  190  forms opening  146 . 
     In the embodiments of  FIGS. 4 and 5 , flange  182  defines surfaces on which a sinter paste (see below) from which sinter joint  180  is formed is deposited, and defines recess  190  in which feedthrough  150  is seated.  FIGS. 4 and 5  illustrate only two embodiments of any number of geometries which may be employed by flange  182  to form recess  190 . According to one embodiment, the width W J  of sinter joint  180  varies depending on a particular type of geometry employed by housing  132  at opening  146 . For example, according to one embodiment, the width W J  of the embodiment illustrated by  FIG. 4  is at least one-fourth the width W F  of feedthrough  150 , while the width W J  of the embodiment illustrated by  FIG. 5  is at least one-half the width W F  of feedthrough  150 . As such, the ratio of the width W J  of sinter joint  180  to the width W F  of feedthrough  150  may vary depending on the geometry of housing  132  about opening  146 . 
       FIGS. 6 through 8  below illustrate and describe embodiments for low-temperature attachment of feedthrough  150  to housing  132  via sintering according to the present disclosure. Sintering is process whereby a solid object is formed from powders, such metal powders (e.g. the powdered bonding material of sinter paste  200 ), by heating, but not melting, the powder. A force is also sometimes applied to compress the powder during the heating process. As opposed to processes where materials are melted, sintering is based on the process of diffusion whereby the atoms in the particles diffuse across particle boundaries as a result of their kinetic energy of random motion, thereby fusing the particles together to form a single, solid piece. Diffusion will occur to some extent in any material above absolute zero, but takes place more rapidly at elevated temperatures. 
     With reference to  FIG. 6 , housing  132  of an implantable medical device is provided, such as implantable medical device  130  of  FIG. 4 , housing  132  including an integral flange  182  having angled and horizontal portions  184 ,  186  forming recess  190  about opening  146 . A sinter paste  200  is applied within recess  190  about a perimeter of opening  146 , with recess  190  serving to hold sinter paste  200  in position. As illustrated, sinter paste  200  is applied with a thickness T P  that is greater than that of the thickness T J  of the finished sinter joint  180  (see  FIG. 4 ). According one embodiment, sinter paste  200  is applied to feedthrough  150  in lieu of housing  132 . In one embodiment, sinter paste  200  is applied to both housing  132  and feedthrough  150 . 
     According to one embodiment, sinter paste  200  includes a biocompatible bonding material in a fine powder or particle form mixed with a binder material. According to one embodiment, as described above, the powdered biocompatible bonding material includes one of gold, platinum, and palladium, or any combination thereof, for example. According to one embodiment, the particle size of the biocompatible bonding material does not exceed 20 μm. According to one embodiment, the particles of biocompatible bonding material are spherical in shape. According to one embodiment, the binder material includes organic solvents, such a butyl terpineol, butyl glycol, and butyl cellusolve, for example. 
     As will be described in greater detail below, the fine particle size enables sintering of the biocompatible bonding material of sinter paste  200  to occur at temperatures much lower than the melting points of the biocompatible bonding materials when in non-powdered form. For example, while the melting point of non-powdered gold is 1,064° C., the temperature at which the sintering effect will occur and cause the gold particles to fuse with one another is well below the 880° C. β-transus temperature of titanium. According to one embodiment, for example, the sintering of gold particles of sinter paste  200  occurs at 350° C. 
     According to the embodiment illustrated by  FIG. 6 , after application of sinter paste  200 , feedthrough  150  is positioned within recess  190  with portions of bottom surface  176  of insulator  152  contacting sinter paste  200 . According to one embodiment, as illustrated, at least the portions of bottom surface  176  contacting sinter paste  200  are provided with a metallized layer  164 . 
     Referring to  FIG. 7 , after feedthrough  150  is positioned on sinter paste  200  so as to cover opening  146 , housing  132  and feedthrough  150  are placed into an oven  210 . According to one embodiment, a support  214  is provided to support a bottom surface  212  of housing  132 , at least in a region of opening  146 , and a weight or anvil  216  is placed on the upper surface  174  of feedthrough  150 . Anvil  216  provides a force F A  which pushes feedthrough  150  toward housing  132  and onto sinter paste  200 , while support  214  provides a counter force F C  to prevent deflection of housing  132  about opening  146  which might otherwise be caused by anvil  216 . 
     At  FIG. 8 , after positioning housing  132  and feedthrough  150  within oven  210 , oven  210  is heated to carry out the sintering process and form finished sinter joint  180  from sinter paste  200 . According to one embodiment, sinter paste  200 , together with housing  132  and feedthrough  150 , are heated to a temperature below the β-transus temperature of the titanium of housing  132  (i.e. about 880° C.) for a desired duration until the finished sinter joint  180  is formed from sinter paste  200  so to arrive at the implantable medical device  130  illustrated by the embodiment of  FIG. 4 . According to one embodiment, oven  210  is heated to a temperature not exceeding 750° C. 
     As described above, although heated to a temperature below the melting point of the powdered bonding material of sinter paste  200  (e.g. the melting point of gold is 1,064° C.), the boundaries of the particles of the powdered bonding material of sinter paste  200  fuse together via the diffusion process to form single solid sinter joint  180 . The smaller the particle size of the bonding material of sinter paste  200 , the lower the temperature and the more quickly the diffusion process will occur. As described above, according to one embodiment, a largest dimension of the particles of the bonding material (e.g. a diameter when the particles are spherical) of sinter paste  200  does not exceed 20 μm. 
     As sinter paste  200  is heated, the binder material is burned off. For example, according to one embodiment, the organic solvents employed as a binder materials for sinter paste  200  (such as those listed above) are burned off at a temperature of approximately 150-160° C. As the binder material is burned off, the volume of sinter paste  200  begins to decrease. As the binder material is burned off, anvil  216  compresses the remaining material of sinter paste  200  into a denser form to ensure that good contact and fusion is made between the powder particles themselves, and between the powder particles and the metallized layer  164  and titanium of housing  132 . 
     Oven  210  has a controlled interior environment  218 . According to one embodiment, in order to enable binder materials of sinter paste  200  to burn off to form the final sinter joint  180 , interior environment  218  is not a vacuum environment. According to one embodiment, in order to prevent oxidation of the titanium of housing  132 , and possibly of metallized layer  164 , which would inhibit the bonding of the binding materials of sinter paste  200  to such surfaces and result in a poor seal therebetween, interior environment  218  is a non-oxygen environment. According to one embodiment, interior environment  218  is one of helium and argon. According to one embodiment, interior environment  218  is one of hydrogen, helium, and argon. 
     Any number of scenarios are envisioned with regard to the heating of oven  210  in order to achieve an optimal sinter joint  180  between housing  132  and feedthrough  150 , wherein heating parameters, such as temperature and duration, may vary depending on a variety of factors, such as the type of bonding materials and binder materials employed by sinter paste  200 , on a thickness with which sinter paste  200  is applied to housing  132 , and on a type of geometry employed about opening  146  (e.g. a shape of recess  190 ), for example. 
     For example, according to one embodiment, where sinter paste  200  employs gold particles as the biocompatible bonding material and organic solvent(s) as the binder material (such as described above), the heating of oven  210  to perform the sintering process includes multiple stages. In a first stage, the temperature is ramped up from an initial temperature of 30° C. to a temperature of 160° C. over a period of 30 minutes. In a second stage, the temperature is maintained at 160° C. for a period of 30 minutes to ensure that the binder materials in sinter paste  200  are completely burned off. In a third stage, the temperature is ramped up from 160° C. to 350° C. and held at 350° C. for a period of 60 minutes to ensure complete fusing (sintering) of the particles of bonding material of sinter paste  200 , in this case gold particles. In a fourth stage, the temperature is ramped down from 350° C. to 30° C. over a period of 60 minutes. Housing  132 , with feedthrough  150  bonded thereto by finished sinter joint  180 , such as illustrated by  FIG. 4 , is then removed from oven  210 . Again, it is noted that any number of heating scenarios may be employed which may include more or fewer than the four steps described by the above example embodiment. 
       FIG. 9  is a flow diagram illustrating a process  300  for hermetically attaching a feedthrough to a housing of an implantable medical device using a sinter joint according to one embodiment of the present disclosure. Process  300  begins at  302  where a titanium housing for an implantable medical device is provided, such as housing  132  of  FIG. 3 . According to one embodiment, housing  132  includes a flange which forms a recess in housing  132 , such as recess  190  of  FIG. 4 . 
     At  304 , a sinter paste is applied about a perimeter of opening  146 , such sinter paste  200  disposed about opening  146  within recess  190  as illustrated by  FIG. 6  for example. According to one embodiment, the sinter paste includes a fine powder of a biocompatible bonding material mixed with a binder material. According to one embodiment, the biocompatible bonding material is one of gold, platinum, palladium, or any alloy combination thereof. 
     At  306 , a feedthrough device, characterized by the absence of a ferrule, is positioned on the sinter paste so as to cover opening  146  in housing  132 , such as ferrule  150  being positioned on sinter paste  200  as illustrated by  FIG. 6 . According to one embodiment, opening  146 , feedthrough  150 , and sinter paste  200  are configured so that the feedthrough  150  overlaps opening  146  such that a finally formed sinter joint has a width at least one-fourth the width of feedthrough  150 , such as illustrated by  FIG. 3 . 
     At  308 , a low-temperature sintering process is performed by heating housing  132 , feedthrough  150 , and sinter paste  200  to a temperature below the β-transus temperature of titanium for a desired duration to form finished sinter joint  180  from sinter paste  200 , such as illustrated and described by  FIG. 8  and the finished sinter joint  180  of  FIG. 4 , for example. According to one embodiment, the low-temperature sintering process is performed in an oven having a controlled environment. According to one embodiment, the low-temperature sintering process includes compressing the sinter paste while being heated. 
     In view of the above, according to the techniques and embodiments of the present disclosure, the attachment of feedthrough  150  to housing  132  using sinter joint  180  eliminates the need for a ferrule (such as ferrule  56  of  FIG. 2 ). By directly integrating feedthrough  150  to housing  132  via sinter joint  180 , as opposed to conventional techniques which integrate the feedthrough to the housing using a ferrule, the shortcomings associated with such a ferrule (e.g. brazed/welded joint, machining requirements, costs) are eliminated. When combined with the use of cermet for conductive elements  154 , feedthrough  150  of the present disclosure provides a complete feedthrough  150  for implantable medical device  130  which is completely devoid of welds and/or brazing. 
     Also, by using a sintering process as described herein to attach feedthrough  150  to the titanium of housing  132  at temperatures below the β-transus temperature of titanium, grain growth within the titanium material of housing  132 , particularly about a perimeter of opening  146 , is greatly reduced relative to conventional techniques which employ high-temperature brazing or welding processes to attach feedthrough devices to housings via a ferrule. As described above, an average grain size of commercially pure titanium employed by housing  132  is initially in the range of about 10-40 μm. 
     According to one embodiment, attaching feedthrough  150  to housing  132  with a sinter joint  180  formed in accordance with the present disclosure results in an average grain size of the titanium of housing  132  proximate to opening  132  that does not exceed 100 μm. As such, according to one embodiment, implantable medical device  130  according to the present disclosure, such as that illustrated by  FIGS. 3-5 , is characterized by a titanium housing  132  having the distinctive structural characteristic imparted by the sintering process described herein of an average grain size not exceeding 100 μm, at least in a region of the housing directly proximate to opening  146 . Such a characteristic is distinctive relative to joints formed by conventional techniques, such as welding and brazing, which result in average grain sizes greatly exceeding 100 μm, such as greater than 300 μm, for example. 
     By reducing the grain growth of the titanium of housing  132 , dimensional distortions of housing  132  are also reduced as compared to conventional techniques, at least to levels whereby dimensions of opening  132  remain within design tolerances after attachment of feedthrough  150 . According to one embodiment, dimensional changes of housing  132  do not exceed 5% relative to initial dimensions. As an example, if opening  146  is a rectangular opening having initial dimensions of 0.020″×0.040″ prior to attachment of feedthrough  150 , the dimensions after attachment using the sintering processes described herein will be within a range 0.019-0.021″×0.038-0.042″. The reduced grain growth also results in the titanium of housing  132 , particularly in the region immediately about the perimeter of housing  132 , becoming less flexible and remaining more rigid as compared to conventional attachment techniques. Reducing dimensional distortions and retaining the rigidity of the titanium about opening  146  reduces the likelihood of a defective or failed connection of feedthrough  150  to housing  132 . 
     Also, because of the low temperatures employed by the sintering process described herein, dimensional changes in housing  132  as a whole are also minimal. The housings of implantable medical devices, such as housing  132  of implantable medical device  130 , are typically formed from two “halves” (one of the halves including opening  146 ) which are later joined to one another, typically by laser welding, to form complete housing  132 . In order to achieve a hermetic seal, the two halves are required to be in close contact with one another during the laser welding process. For example, the dimensional profiles of the two halves of the housing are typically required to be held within a tolerance of +/−0.004″. The low-temperatures associated with the sintering process described herein ensure that the dimensional profiles of the housing halves remain within required tolerances. 
     Finally, because sintering is a low-temperature process whereby the boundaries of the powdered metal particles fuse together to form a solid body in the absence of melting, a sinter joint typically includes gaps or voids within the solid body. In contrast, in joints formed by welding or brazing the bonding material is completely melted and reflows such that the resulting solid joint has virtually no voids or gaps. As a result, whereas a joint formed by brazing or welding has a density of greater than 99% (i.e. the joint comprises greater than 99% bonding material), a joint formed by sintering as described herein has a density of 90-99% (i.e. the joint comprises 90-99% bonding material, the remainder being gaps or voids). 
     For example, a brazed joint of pure gold would have a density of greater than 19.1 g/cm 3  (i.e. the density of pure being 19.3 g/cm 3 ), while a sinter joint of pure gold would have a density in the range of about 17.4-19.1 g/cm 3 . As such, according to one embodiment, implantable medical device  130  according to the present disclosure, such as that illustrated by  FIGS. 3-5 , is characterized by a distinctive structural characteristic imparted by the sintering process of sinter joint  180  having a 90-99% density of bonding material (e.g. gold, gold alloy). It is noted that even though a sinter joint is less dense than a welded or brazed joint, sinter joint  180  still provides a hermetical seal between the feedthrough  150  and housing  132 . 
     Although specific embodiments have been illustrated and described herein, it will be appreciated by those of ordinary skill in the art that a variety of alternate and/or equivalent implementations may be substituted for the specific embodiments shown and described without departing from the scope of the present invention. This application is intended to cover any adaptations or variations of the specific embodiments discussed herein. Therefore, it is intended that this invention be limited only by the claims and the equivalents thereof.