DMLS orthopedic intramedullary device and manufacturing method

Abstract

The present invention relates to DMLS orthopedic intramedullary devices and methods of manufacture. An orthopedic device for internal fixation of bones, such as an intramedullary nail, and a method of manufacture thereof. The orthopedic device can be formed from medical grade powders by an additive manufacturing process. The forming process can include heat treating the additively manufactured parts and processing the heat treated additively manufactured parts to form the orthopedic device. In addition, the orthopedic device can be formed to include an internal sensor probe channel that extends within at least a portion of the wall of the device but does not protrude through the exterior of the wall. Embodiments also include a dynamic intramedullary nail that accommodates adjustment of the relative axial position of one or more sections of the orthopedic device. The device may include features in an internal region of the orthopedic device that can change the elastic modulus of the orthopedic device.

Description

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS

[0002] The present invention is a divisional application of Chinese patent application No. 2015800313579. This application claims the benefit of U.S. Provisional Patent Application Serial No. 61 / 978,804 filed on April 11, 2014 and U.S. Provisional Patent Application Serial No. 61 / 978,806 filed on April 11, 2014, both of which are incorporated herein by reference in their entirety. Technical Field

[0003] The present invention relates generally to implants for use in orthopedic surgery or procedures, and more particularly, but not exclusively, to orthopedic intramedullary devices, such as orthopedic intramedullary nails, for internal fixation of bones, and methods of making the same. Background Art

[0004] Orthopedic fixation devices can be used, for example, to stabilize injuries, support fractures, fuse joints, and / or correct deformities. Orthopedic fixation devices can be attached permanently or temporarily, and can be attached to bones at various locations, including implanted in a tube or other cavity of the bone, implanted under soft tissue and attached to the outer surface of the bone, or disposed externally and attached by fasteners such as screws, pins, and / or wires. Some orthopedic fixation devices allow the position and / or orientation of two or more bone fragments or two or more bones to be adjusted relative to each other. Orthopedic fixation devices are typically machined or molded from isotropic materials, such as metals including, for example, titanium, titanium alloys, stainless steel, cobalt-chromium alloys, and tantalum.

[0005] Furthermore, the primary function of an intramedullary (IM) nail is to stabilize the fracture fragments and thereby enable load transfer across the fracture site while maintaining the anatomical alignment of the bone. Although there are a large number of different commercially available intramedullary nails on the market, there are no universal guidelines stating the conditions under which each nail will perform at its best in a given situation. Furthermore, the optimal implant stiffness level is a topic of some debate, and the mechanisms underlying the interaction between the local mechanical environment and fracture healing are generally unknown.

[0006] In addition, in terms of torsion and bending, the effect of the fixed stiffness of change on fracture healing can provide insights into the pathogenesis and ideal treatment of fracture and non-union. However, at least due to cost control reasons, similar implants are used for both simple and complex fractures. Therefore, finding a relatively optimal solution in terms of axial bending and torsional stiffness closer to bone rather than titanium or stainless steel may accelerate the fracture healing of a specific type of fracture.

[0007] There remains a need to provide improved orthopedic intramedullary devices for internal fixation of bones and methods of making the same. The present invention satisfies this need and provides other benefits and advantages in a novel and unobvious manner. Summary of the invention

[0008] One aspect of the present invention is a method for manufacturing an orthopedic device, comprising forming an additively manufactured orthopedic component from a medical grade powder and through an additive manufacturing process. The method also includes heat treating the additively manufactured orthopedic component and processing the heat treated additively manufactured orthopedic component to form an orthopedic device.

[0009] Another aspect of the present invention is an intramedullary nail comprising a wall comprising one or more laser sintered layers of medical grade powder. The wall has an exterior and an interior, the interior generally defining an interior cannulation region of the intramedullary nail. The intramedullary nail also includes an interior channel for receiving a miniaturized sensor probe extending into at least a portion of the wall. Additionally, the interior sensor probe channel does not extend through the exterior of the wall.

[0010] In addition, one aspect of the present invention is an intramedullary nail having a first section coupled to a second section by a telescoping portion. The telescoping portion has an outer diameter sized to be slidably received in an interior region of at least one of the first and second sections to accommodate adjustment of the relative axial position of the first and second sections. In addition, the first and second sections are structures for implantation in a bone. The intramedullary nail also includes a mechanical actuator adapted to provide a biasing force to bias the relative axial position of the first and second sections.

[0011] Another aspect of the present invention is an intramedullary nail comprising a wall having an exterior and an interior, the interior generally defining an interior region of the intramedullary nail. The intramedullary nail further comprises a first screw hole and a second screw hole, the first and second screw holes extending at least through the exterior of the wall. In addition, the intramedullary nail comprises one or more protrusions in the wall between the first and second screw holes, the one or more protrusions not extending at least through the exterior of the wall. In addition, the one or more protrusions are configured to change the torsion and bending modulus of the intramedullary nail.

[0012] Another aspect of the present invention is an intramedullary nail having a first section with a wall that substantially defines an interior region of the first portion. The intramedullary nail also includes a second section coupled to the interior section, the interior section being dimensioned to be laterally displaced in at least a portion of the interior region of the first section. In addition, the interior section is selectively removable from the first section by a locking screw to selectively change the mechanical properties of the intramedullary nail. BRIEF DESCRIPTION OF THE DRAWINGS

[0013] The description herein refers to the drawings, wherein like reference numerals refer to like parts throughout the several views.

[0014] Figure 1A An isometric view of a three-dimensional (3D) model of a hypothetical construction program with a vertical stack of 143 pegs is shown.

[0015] Figure 1B A top view of a three-dimensional (3D) model of a hypothetical construction program with a vertical stack of 143 pegs is shown.

[0016] Figure 1C A side view of a three-dimensional (3D) model of a hypothetical construction procedure with a vertical stack of 143 pegs is shown.

[0017] Figure 2A Shown is a microscopic view of the distal axis of a direct metal laser sintering (DMLS) intramedullary nail with 30% porosity.

[0018] Figure 2B Shown is the nail fracture site at the distal position of the intramedullary nail after a 4-point bending fatigue test.

[0019] Figure 3A An ALM nail is shown manufactured with minimal post-machining requirements, ie, all design features turned on in the CAD file.

[0020] Figure 3B An ALM pin manufactured with all design features closed in the CAD file is shown along with the support structure used to hold the part in the chuck of the CNC machine.

[0021] Figure 3C An ALM nail is shown manufactured with the design features in the CAD file turned off at the distal end.

[0022] Figure 4 Shown is the time to build a nail on an SLM500, which scales approximately linearly with the size (length) of the nail.

[0023] Figure 5 A comparison of example build times for 16 test rods is shown, including an interpolation of the time and cost to build a 280 millimeter (mm) high intramedullary nail using different machines.

[0024] Fig. 6A A perspective view of manufacturing an intramedullary nail using a vertical orientation is shown.

[0025] Figure 6B A perspective view of manufacturing an intramedullary nail using a horizontal orientation is shown.

[0026] Fig. 7A A three-dimensional (3D) CAD intramedullary nail model including a tapered inner wall section in a cross-sectional view is shown.

[0027] Figure 7BA three-dimensional (3D) CAD intramedullary nail model is shown in a partial perspective view and includes a porous or channel internal structure in the wall of the intramedullary nail.

[0028] Figure 7C A three-dimensional (3D) CAD intramedullary nail model including a removable inner section is shown in perspective, partial cross-sectional view.

[0029] Fig.7D A three-dimensional (3D) CAD intramedullary nail model including an inner grooved section is shown in partial cross-section.

[0030] Figure 8 Examples of optimal cross-sectional geometries for intramedullary nails are shown.

[0031] Fig. 9 Schematic diagram showing the layer-by-layer laser sintering additive manufacturing process parameters reproduced from a build performed on a Renishaw SLM250.

[0032] Fig. 10A A cross section of an intramedullary nail is shown highlighting the wiring area in the wall segment associated with the double scanning strategy.

[0033] Fig. 10B Unidirectional X and Y scans are shown in association with a dual scanning strategy for remelting a constructed laser sintered layer.

[0034] Fig. 10C Multi-directional X and Y scanning associated with a dual scanning strategy for remelting a constructed laser sintered layer is shown.

[0035] Fig. 10D Parameters of the sintered layers of the remelting structure using the double scanning strategy are shown.

[0036] Fig.11A A cross section of an intramedullary nail is shown, highlighting a wall segment of the intramedullary nail associated with an X and Y alternating line laser raster of a double scanning strategy for remelting the constructed laser sintered layer.

[0037] Fig. 11B An X and Y alternating line laser raster associated with a dual scanning strategy for remelting a constructed laser sintered layer is shown.

[0038] Fig. 11C A cross section of an intramedullary nail is shown, highlighting a wall portion of the intramedullary nail associated with a circumferential laser raster of a double scanning strategy for remelting the constructed laser sintered layer.

[0039] Fig.11D A circumferential laser grating associated with a dual scanning strategy for remelting a constructed laser sintered layer is shown.

[0040] Fig.12 The hammer head stress distribution of a Ti-64 Additive Layer Machined (“ALM”) part is shown.

[0041] Fig.13 A three-dimensional model of an ALM test section attempting to clone the distal end of a standard intramedullary nail is shown.

[0042] Fig.14 An example of the effect of laser power on porosity as a function of position along the ALM sample is shown.

[0043] Fig.15 Shown is an SEM image taken from the fracture surface of an ALM part sintered to 400 watt (W) laser.

[0044] Fig.16 A table depicting the results of four-point bending tests conducted at 5 Hertz (Hz) is shown. Machined Ti6-4 nails were cut to similar lengths and tested using the 300 to 3000 Newton (N) method to validate the use of the test rig with short nails. The nails survived 10 6 The ALM samples were tested using the same method but with load conditions ranging from 200 to 2000 Newtons (N).

[0045] Fig.17 The effect of laser power on the four-point bending fatigue properties of a Ti-64ALM part performed at 5 Hertz (Hz) under a step load of 2000:200 Newton (N) is shown.

[0046] Fig.18 A list of exemplary machine suppliers for laser sintering and electron beam melting of Ti-64 material is provided.

[0047] Fig.19A A 2D diagram of the simplified test specimen geometry is shown.

[0048] Fig.19B Photographs showing post-machining / heat treatment of a cannulated Ti-64 specimen.

[0049] Fig. 20 Four-point bending fatigue performance of laser sintered and electron beam melted specimens as a function of machine vendor is shown. Tests were performed at 200-2000 N and 5 Hertz (Hz).

[0050] Fig.21A and Fig.21B Microscopic sections of EOS samples at different magnifications are shown and etched with Kroll's reagent to visualize the particle structure.

[0051] Fig. 22 Four-point bending fatigue performance of HIPPED laser sintered and electron beam melted specimens as a function of machine vendor is shown. Tests were performed at 300 to 3000 N and 5 Hertz (Hz).

[0052] Fig.23A Shown is the crystal structure of Grade 23 titanium produced as a fabricated part on an EOS M280 machine prior to hot isostatic pressing treatment.

[0053] Fig. 23B shows the components after they have been subjected to hot isostatic pressing. Fig.21A Grade 23 titanium produces a crystalline structure for the components.

[0054] Fig.24 The generally ordered structure of a conventional lamellar-type Ti-6Al-4V titanium wrought alloy heated to the beta transus temperature of 1005°C and cooled at a rate of 100°C min-1 to 700°C, stabilized for 2 hours, and then cooled in an oven is shown.

[0055] Fig.25 Tables are provided showing the surface finish or roughness of finished parts that have undergone post-processing.

[0056] Fig.26 The average measured thickness of the alpha flakes in the ALM samples studied is presented.

[0057] Fig. 27 The key parameters determining the fatigue properties of platelet-type titanium alloys are provided. (D) = particle size, (t) = width of the platelet, (d) = size of the colony of parallel platelets.

[0058] Fig.28 Shown are the surface finish and four-point bending fatigue properties of HIPPED laser sintered and electron beam melted specimens as a function of machine supplier. Tests were performed at 5 Hertz (Hz) from 400 to 4000 Newtons.

[0059] Fig.29 Load extension curves for heat treated and non-heat treated ALM parts are shown. Points F and J represent the performance of the non-heat treated parts. These parts are strong but brittle.

[0060] Fig.30 Strength (UTS) and ductility (% elongation) data captured from as-forged and ALM Ti-64 parts are shown.

[0061] Fig.31 Calculations are shown for producing 100 fully dense intramedullary nails using single, dual, and quad lasers at normal scan speed.

[0062] Fig.32Calculations for producing 100 fully dense intramedullary nails at ultrascan speeds are shown.

[0063] Fig.33A An exemplary breakdown of production costs ($) per step for manufacturing 100 intramedullary nails using standard scanning conditions is shown. Assumptions: - Cost of running the ALM machine = $97 per hour. Cost of implant grade powder = $255 per Kg. 96 hour build time to manufacture 100 nails.

[0064] Fig.33B Shows Fig.33A The production costs shown in , as a percentage cost of each manufacturing step.

[0065] Fig.34A shows the use ratio with Figure 31-32 The example related scanning conditions shown are an exemplary breakdown of the production costs for manufacturing 100 intramedullary nails.

[0066] Fig.34B Shows Fig.34A The production costs shown in , as a percentage cost of each manufacturing step.

[0067] Fig.35 The proximal end of the Trigen Meta tibial nail is shown and the concept of boundary layer scanning is highlighted.

[0068] FIG. 36A shows a model of an orthopedic intramedullary nail after boundary scanning and adjusted in size to account for size shrinkage.

[0069] Fig.36B A model of an orthopedic intramedullary nail after HIPPING is shown, indicating a fully dense part that meets the required CAD file part specifications.

[0070] Fig.37A A photograph showing a powder core produced by 5% boundary scan "In situ shelling" is shown.

[0071] Fig.37B A CAD file is shown highlighting the distal extrusion section used to prevent residual powder from escaping the core after removal from the build plate.

[0072] Fig.38 An optical image taken from the polished surface of the wall segment of the distal tip after heat treatment is shown. The porosity and average pore size were measured to be approximately 0.25% and 3.4 microns, respectively.

[0073] FIG. 39A shows a model of the orthopedic intramedullary nail after ultralaser scanning and adjusted in size to account for size shrinkage.

[0074] Fig.39B A model of an orthopedic intramedullary nail after HIPPING is shown and indicates a fully dense part that meets the required CAD file part specifications.

[0075] Fig.40 Four-point bending fatigue performance of ALM parts produced by SLM solutions subjected to variable scanning strategies (full skin, no core, full skin, partial core) is shown.

[0076] Fig.41 A sagittal section through a HIPPING oven is shown, highlighting the exposure of the ALM portion to silver vapor.

[0077] Fig.42 A schematic diagram of the bead blasting procedure is shown, highlighting the deposition of silver onto a workpiece during a finishing operation.

[0078] Fig.43A A three-dimensional (3D) model of an intramedullary nail having a built-in channel extending generally parallel to the cannula to temporarily house a sensor probe is shown.

[0079] Fig.43B A microCT image of a nail with a built-in channel (1.5 mm diameter) is shown, highlighting the absence of residual powder in the channel after construction.

[0080] Fig.44 A three-dimensional (3D) model of a removable sensor probe that may be designed to operate within the built-in sensor probe channel of an intramedullary nail shown in FIG. 43 is shown.

[0081] Fig.45A and Fig.45B Shown is a model of an intramedullary nail having an open channel formed in the proximal end and outer surface of the intramedullary nail suitable for conventional manufacturing techniques.

[0082] Fig.46A An end view of an exemplary geometry of an internal sensor probe channel that can be produced using additive manufacturing is shown.

[0083] Fig.46B An isometric view of an exemplary geometry of an internal sensor probe channel that can be produced using additive manufacturing is shown.

[0084] Fig.47 An additively manufactured nail is shown, including an internal sensor probe channel.

[0085] Fig.48 A perspective view of the proximal end of an intramedullary nail constructed by additive manufacturing and including a built-in probe channel adapted to receive insertion of a removable sensor probe is shown.

[0086] Fig.49 A perspective view of the distal end of a dynamic intramedullary nail according to an exemplary embodiment of the present invention is shown.

[0087] Fig.50A , Fig.50B and Fig.50C A dynamic intramedullary nail having telescoping portions at the distal, middle, and proximal regions of the intramedullary nail, respectively, is shown.

[0088] Fig.51 Examples of linear and nonlinear spring force-displacement relationships are shown.

[0089] Fig.52A and Fig.52B A schematic diagram of the degradation of a resorbable polymer of a dynamized, actuated loaded intramedullary nail in a long bone fracture is provided.

[0090] Fig.53A and Fig.53B A dynamic intramedullary nail having a telescoping portion configured to provide unidirectional and bidirectional translation of the intramedullary nail, respectively, is shown.

[0091] Fig.54 and Fig.55 A dynamic intramedullary nail having a telescoping portion comprising a protrusion in the form of a pin is shown.

[0092] Fig.56 The effect of wall thickness on mechanical stiffness measurements of intramedullary nails is shown.

[0093] Fig.57 Shown are perspective views of examples of three bushings configured for use with an intramedullary nail to locally generate cyclic loading at a fracture site.

[0094] Fig.58A An external longitudinal view of the end of an intramedullary nail with an activatable shape memory sleeve or collar is shown.

[0095] Fig.58B and Fig.58C Shows Fig.58A A perspective view of the interior of an intramedullary nail is shown, which includes an activatable shape memory sleeve or ring in active and inactive states, respectively.

[0096] Fig.59 A schematic diagram of portions of a dynamic intramedullary nail including a pair of opposing encapsulated biasing elements is shown.

[0097] Fig.60 Examples of standard circular cross-section intramedullary nails and cross-sectional geometries of intramedullary nails that can reduce bending stiffness in the anterior-posterior plane while relatively maintaining stiffness in the orthogonal medio-lateral plane are shown.

[0098] Fig.61A and Fig.61B Torsional and bending stiffness data based on a standard 10 mm outer diameter Trigen Meta tibial nail and a 10 mm outer diameter Trigen Meta tibial nail having grooves of varying sizes on the outer surface of the nail are shown.

[0099] Fig.62 A three-dimensional (3D) CAD intramedullary nail model is shown in cross-section, including a tapered inner wall at the mid-section of the nail.

[0100] Fig.63 A table indicating the theoretical bending and torsional stiffness of a mid-section of an intramedullary nail having a tapered inner wall is shown.

[0101] Fig.64A A partial perspective view and a cross-sectional view of a three-dimensional (3D) CAD model of an intramedullary nail provided with circumferentially arranged internal grooves in an inner wall section of the nail are shown.

[0102] Fig.64B Shows Fig.64A A cross-sectional view of a portion of the middle section of an intramedullary nail is shown.

[0103] Fig.65 A three-dimensional (3D) CAD model of an intramedullary nail comprising an ordered porous or channel internal structure in the wall of the intramedullary wall is shown in partial cross-sectional view.

[0104] Fig.66 The image taken from the middle section of the intramedullary nail is shown. Fig.65 An end view of a portion of an intramedullary nail is shown.

[0105] Fig.67 The standard intramedullary nail and Fig.66 and 65 A comparison of the theoretical bending and torsional stiffness of intramedullary nails having a porous or channeled internal structure is shown.

[0106] Fig.68 Shown are elastic modulus versus density plots providing data for cancellous bone, collagen, Ti 40% porosity, cortical bone, TI-6AI-4V 40% porosity, Ti alloy, CO-Cr alloy, steel, and Ti-6A-4V.

[0107] Fig.69 A perspective, partially cut-away view of a portion of a three-dimensional (3D) model of an intramedullary nail including a removable inner section is shown.

[0108] Fig.70 Shown is a standard Trigen Meta tibial nail and a Fig.69Table showing comparison between the bending and torsional stiffness of nails with similar removable inner sections.

[0109] The foregoing summary of the invention and the following detailed description of certain embodiments of the invention will be better understood when read in conjunction with the accompanying drawings. For the purpose of illustrating the invention, certain embodiments are shown in the accompanying drawings. However, it should be understood that the invention is not limited to the arrangements and tools shown in the accompanying drawings. DETAILED DESCRIPTION

[0110] In order to promote the understanding of the principles of the present invention, reference will now be made to the embodiments shown in the accompanying drawings, and specific language will be used to describe these embodiments. However, it should be understood that the scope of the present invention is not limited thereto. It will generally be appreciated by those skilled in the art to which the present invention relates that any changes and further modifications in the described embodiments, as well as any further application of the principles of the present invention described herein, will be appreciated. The following description and illustration of the non-limiting forms and embodiments of the present invention are exemplary in nature, and it should be understood that the description and illustration related thereto are by no means intended to limit the present invention and / or its application and use disclosed herein.

[0111] Orthopedic devices require certain material properties and / or tolerances for optimal manufacturing and performance under stress loading conditions within the human body. For fixation devices, such as intramedullary (IM) nails, such properties or characteristics may include four-point bending fatigue, flexural modulus, torsional stiffness, tensile strength / ductility / yield strength, porosity, surface finish, and geometric tolerance / part accuracy. While conventional forged / machined Ti-64 nails may meet the current required standards, with the advent of rapid manufacturing technologies (RMT), there is an opportunity to reduce the overall cost of raw manufacturing and goods relatively significantly. In addition, global market sales of implantable devices manufactured using titanium are estimated to reach $26 billion by 2020, which emphasizes the need to change the manufacturing process to meet the expected volumes. In addition, additive manufacturing has the advantage of providing near-net-shape products / components to multiple markets without having to rely on a highly skilled workforce. Additionally, given its design freedom, the use of additive manufacturing in the design and manufacture of implant devices such as intramedullary nails may open up the possibility of developing implants specific to the patient's characteristics, including, for example, the patient's age, bone quality, and injury type, among other characteristics.

[0112] RMT techniques include, but are not limited to, direct metal fabrication (DMF), direct metal laser sintering (DMLS), electron beam welding (EBM), and solid free-form manufacturing. These techniques have been used in a variety of industries, including orthopedics for reconstruction, trauma, and rehabilitation devices. In general, DMLS can use three-dimensional (3D) computer-aided design (CAD) models that can be created using programs such as Materialise. ®Magics ® , to produce a three-dimensional metal sintering model produced layer by layer by irradiating metal powder with a laser. For example, Figures 1A-1C Shown are Magic ® Isometric, top, and side views of a three-dimensional (3D) model of a hypothetical construction program created by the program, with a vertical stack of 143 pegs on a 25 cm by 25 cm (x by y) construction board.

[0113] However, the use of DMLS can create a number of issues related to material properties and functionality that cannot be tolerated in orthopedic devices, including porosity, part tolerances, part design, and surface finish, which require additional post-processing, each of which will be summarized in turn below.

[0114] (A) Porosity: Suboptimal laser beam optics can create porous regions within the material, which leads to poor material properties and subsequent poor / reduced performance. These porosity issues are often primarily attributed to the entrapment of unsintered powder within the fabricated component and residual gases such as argon and oxygen during sintering. For example, Figure 2A The distal shaft of a DMLS intramedullary nail with 30% porosity at the distal axis is shown. Figure 2B The associated nail fracture sites at the distal position of the intramedullary nail as a result of a 4-point bending fatigue test are shown.

[0115] (B) Part Tolerances: A suboptimal laser beam (power) can produce a component with specifications outside the required tolerances. For example, for an intramedullary nail, such areas outside the required tolerances that can be produced by a suboptimal laser beam include the inner / outer portions of the nail and the area associated with the internal screw diameter.

[0116] (C) Component Design: Given the complex geometry of trauma fixation devices such as intramedullary nails, it is critical to determine which design features should be opened during the laser sintering stage. Likewise, identifying the appropriate support structure design required to hold the component during post-processing is also critical to the quality of the final component. Figure 3A A section is shown with all design features turned on during the build phase. Design features at the far end of the component are particularly sensitive to fatigue stresses given the thin walls. Components are typically constructed in a vertical orientation, and therefore, the microstructure around the transverse screw holes is particularly susceptible to defects due to the nature of heat dissipation in the component. Therefore, despite requiring minimal post-machining, this component design does not produce satisfactory mechanical fatigue performance. Figure 3BA more conservative design is depicted, as all design features (lateral screw holes, keyways, etc.) have been closed during the build phase / The component is also equipped with a protrusion at the proximal end to secure the part in the chuck of the CNC machine. Although this design yields satisfactory mechanical properties, the post-machining phase is very intensive. Figure 3C The implant design is depicted to provide a balance between post-machining needs and desired mechanical properties. The distal features are closed at the most vulnerable part of the nail, which reduces post-machining time but addresses the risk of early fatigue failure.

[0117] (D) Additional processing: Currently, the optimal post-DMLS processing to impart mechanical properties is often unknown or suboptimal. Techniques are known in the art, such as hot isostatic pressing (HIP), in which the implant is subjected to elevated temperature and isostatic gas pressure to consolidate and reduce porosity within the material; shot peening, in which the implant is bombarded to produce plastic deformation and remelting the sintered layer to reduce the porosity of the growing part are other strategies aimed at improving component performance.

[0118] In addition, the use of DMLS can create a number of issues related to commodity costs and productivity. For example, in at least some cases, depending on the grade of Ti-6AL4V powder purchased, such as plasma rotating electrode process (PREP), gas atomization, and ELI grade 23, the cost of metal powder can be about $200-400 / kg. In addition, the machine cost / build time can be about $97 / hour, depending on the size of the manufacturing unit, capital equipment and depreciation levels, and staff. In addition, a high degree of nailing can be used to obtain a part throughput of 100 nails per component, which can take about 109 hours to complete (65.4 minutes per nail) for a subtractive machining operation performed in a medical device manufacturing unit, compared to, for example, 600 nails per day (2.4 minutes per nail). A laser sintering machine with a larger footprint and equipped with 4 scanners (e.g., an SLM 500 quad scanner) can be used to increase the efficiency of the additive manufacturing process. 800 pins (dimensions 20cm long x 10mm OD) can be built in 160 hours, which is equivalent to 12 minutes per pin using a boundary scan "in-situ shelling" strategy. If running to full capacity and using a boundary scan strategy, the SLM500 is capable of producing about 40,000 pins per year. If the supply of atomized powders at a price of £30 / Kg can be used in preference to material supplied from powder suppliers (usually $250-400 / Kg), the average cost per pin can be close to $120 per part. Figure 4 The time to build a nail on the SLM500 is shown, and the SLM500 scales approximately linearly with the size (length) of the nail. For example, a 320 mm nail requires an average per-nail build time of 25 minutes, Figure 4 .

[0119] The cost of metal powders is often controlled by the supplier, who may charge a fee for a limited number of suppliers in certain markets. For example, European supply companies, more specifically UK suppliers, often offer substantially cheaper and environmentally friendly powders compared to existing titanium production methods (e.g., energy-intensive gas atomization and the toxic Kroll process), which constitute an expensive and labor-intensive four-step process. Such suppliers can take rutile and convert it directly into powdered titanium using electrolysis, which is cost-effective and therefore often necessary for the supply chain. Low-cost titanium powders can be used in a variety of new applications where previously the metal was too expensive to use for mass production of low-value items. For example, gas atomized powders directly from bar stock are a potential route to reduce the cost of Ti-64 power to approximately £30 / Kg.

[0120] Figure 5 An example of a comparison of build times for 16 test bars using different machines identified as Arcam S12, Concept M2, EOS M270, Realizer, Renishaw AM250 and SLM Solutions 280HL, including insertion time and cost to build a 280 millimeter (mm) high intramedullary nail is shown. The build time of a part can be affected by a number of interrelated variables including, for example, DMLS machine specifications, as well as operating costs such as gas, electricity and capital equipment and other costs such as Figure 5 While there may be relatively significant scope to reduce the cost of part manufacturing, it is not necessarily comparable to the cost of machining those parts, which eliminates the need for other cost reduction strategies such as part design, scan speed, and scan pattern.

[0121] The present invention provides an optimal DMLS manufacturing route for orthopedic devices with material properties matching those of wrought / cast / machined titanium components.

[0122] In one form of the invention, a method for manufacturing an elongated orthopedic device by direct laser sintering is provided, comprising the steps of: a) generating a virtual three-dimensional (3D) elongated device model; b) using a laser power of at least 300 Watts (W) and using a powder of at least Grade 5 quality, such as TiAl6v4 powder, to manufacture the elongated device in appropriate build directions according to the three-dimensional (3D) model by direct metal sintering; c) subjecting the elongated device to hot isostatic pressing (HIP) using a temperature of at least 1000 degrees, wherein the cooling rate is between 0.24 and 72 degrees Celsius min-1; d) machining and polishing the HIP treated elongated device; and e) wherein the mechanical properties are equivalent to the four-point bending properties of wrought titanium.

[0123] In another form of the invention, an orthopedic implant, such as an intramedullary nail, is manufactured by the following steps and processes:

[0124] (A) CAD file creation: Upload the appropriate file type (e.g., .stl formatted file) in a manufacturing-friendly orientation to a three-dimensional (3D) software provider such as Materialise ® Magics ® Such files may include, for example, non-supportive vertical building structures of intramedullary (IM) nails, as well as other components or devices.

[0125] (B) Build Orientation: Parts can be built in multiple build orientations from 0 to 90 degrees, which will produce anisotropic physical properties that are sensitive to mechanical loads. They can also be built with or without custom designed support structures to assist in post-machining processes. Figure 3B In addition, building components such as intramedullary nails vertically using an Additive Layer Machine (“ALM”) equipped with a soft recoater and without any support structures can reduce the burden of post-processing. Fig. 6A and Figure 6B shows the use of vertical orientation ( Fig. 6A ) and horizontal orientation ( Figure 6B ) is a three-dimensional (3D) CAD file for the fabrication of a prominent intramedullary nail, wherein the vertical and horizontal orientations are suitable for an ALM with soft and hard recoats. Fig. 6A The build section is shown in a vertical configuration / orientation and can be made more economical by designing a support structure that allows stacking of components. Figure 6B The build parts are shown in configurations / orientations at angles between 0 and 90 degrees and may help reduce their anisotropic behavior. However, a smaller number of devices may be manufactured in a single build run.

[0126] (C) Design Optimization: RMT can be used to grant considerable design freedom regarding the internal and external geometry of the intramedullary nail (Figure 6) to modify nail properties, such as torsional / bending stiffness, which are difficult to achieve using conventional manufacturing methods. Specifically, Figures 7A-7D A three-dimensional (3D) CAD intramedullary nail model is shown in cross-section and highlights the tapered wall segments ( Fig. 7A )、Porous internal structure( Figure 7B ), removable internal section ( Figure 7C ) and the inner grooved section ( Fig.7D ).in addition, Figure 8Examples of optimal cross-sectional geometries for intramedullary nails are shown, where: A = Solid Schneider, B = Diamond, C = Sampson Fluted, D = Kuntscher, E = Rush, F = Ender, G = Mondy, H = Halloran, I = Huckstep, J = AO / ASIF, K = Grosse-Kempf, L = Russell Taylor, M = Trigen. At least some of these nail geometries may have cross-sections designed to reduce stiffness or reduce intramedullary pressure. In addition, the diversity of these nails ranges from the Kuntscher nail (D), which may be a tight fit, reamed, non-locking, to the Universal nail (K), which includes the use of interlocking screws.

[0127] RMT can also be used to produce patient-matched implants by optimizing the curvature of the intramedullary nail. This can avoid a mismatch in the radius of curvature between the intramedullary nail and the bone, particularly the distal femur, which could otherwise lead to anterior cortical perforation. For example, the radius of curvature of the femur is estimated to be 120 cm (+ / - 36 cm). However, femoral nail designs typically have smaller curvatures, with radii ranging from 186 to 300 cm. In addition, the nail can also be designed to fit each individual fracture. According to such an embodiment, a computer model of each individual fracture can be created, which can then be used to test different fixation strategies in order to select a system that will create a specific mechanical environment for the assumed load-bearing requirements.

[0128] (D) Three-dimensional (3D) printer selection: Intramedullary nails can be constructed using a variety of commercial machines from vendors such as SLM Solutions, Renishaw, Realizer, EOS, Concept Laser, and Arcam. The relative merits of each technology are generally based on: (a) machine productivity (i.e., chamber size (along the x, y, and z axes), scan speed, and number of lasers); (b) part quality (i.e., accuracy, surface finish, tolerance / resolution); and (c) capital and operating costs (e.g., gas and electricity consumption).

[0129] (E) Laser Sintering: Hard recoaters, such as the EOS M270 / M280 / M290 from EOS can produce parts with excellent mechanical properties and reduced porosity, considering that any weakly bound, partially sintered material is more likely to be removed at each build layer. Soft recoaters can produce parts that may also be contaminated with silicone blade fragments, which requires investigation to meet regulations. Soft recoater blades may wear after a build, which adds additional cost to the manufacturing process. Hard recoater blades can be made of high-speed steel, and debris released from these arms into the part has been observed to cause fewer problems than soft recoater blades. Hard recoater blades are also more economical for powder use. Modern laser sintering machines provided by SLM Realizer can produce a focused beam spot size of 30 microns, which can produce parts with excellent grain structure and resolution, enabling novel design features such as built-in channels.

[0130] (F) Powder Specifications: Medical grade Ti-64 powders are available in a number of different styles, depending on the end application and the choice of three-dimensional (3D) printer. Grade 5 gas or plasma atomized powders are commonly used in laser sintering and may have a particle size range of 15 to 45 microns (μm) or 20 to 63 microns (μm), and are typically supplied at a cost of £150 / kg. Grade 23 ELI powders may be gas atomized or centrifugal PREP powders with a particle size range of between 45 and 100 microns (μm) and are typically supplied at a cost of £250 / kg, and may contain reduced levels of oxygen, nitrogen, carbon and / or iron. The decision to switch to virgin powder for subsequent builds, or to utilize unused powder from previous builds, may be determined from QA testing of the built parts and powder bed. Obviously, building parts from unused powder left over from previous builds will reduce the cost burden and need to purchase several tons of powder to cover the manufacture of a large number of intramedullary nails (i.e., 10,000 or more).

[0131] (G) Selection of Build Parameters: In the production of intramedullary nails, the build parameters selected for additive manufacturing technologies including selective layer melting (SLM), laser sintering, and electron beam processing, among other technologies, may be unique to the brand or manufacturer of the equipment used to perform the additive manufacturing. For example, laser sintering or electron beam processing of Ti-64 may be unique to the machine tool supplier. In addition, Fig. 9 A schematic diagram of the build parameters that can be used with a Renishaw SLM250 to produce an intramedullary nail is provided. Specifically, Fig. 9Schematic representation of the laser sintering additive manufacturing process parameters reproduced layer by layer from a build performed on a Renishaw SLM250 with a laser power between 50 Watts (W) and 280 Watts (W), a spot distance between 30 and 90 micrometers (μm), a line distance of 65 μm, a layer thickness of 50 μm and an exposure (Ex) of 50 to 500 microseconds (μs).

[0132] Regarding scanning strategies, while some additive manufacturing techniques, such as Selective Layer Melting (SLM), are able to produce fully dense materials, it may be necessary or beneficial to reduce partial porosity using one or more of the following scanning strategies:

[0133] (1) Layer remelting: Figures 10A-10D As shown, remelting of the laser sintered layer can help reduce the porosity of the growing part using a dual scanning strategy. In addition, the variable focus optics in some systems, such as the 1 kilowatt (kW) SLM Renishaw SLM250 system, can enable the use of high laser power at a relatively slow speed, with a relatively large spot size in the center of the layer to compensate for high heat losses. In addition, high laser power at a relatively high speed can be utilized at the surface area (i.e., the boundary of the layer) in order to achieve high surface quality. Fig. 10A A cross section of an intramedullary nail is shown, highlighting the wiring area in the wall section of the intramedullary nail associated with the double scanning strategy. Fig. 10B shows unidirectional X and Y scans (i.e., the directions are parallel to each other), and Fig. 10C Multi-directional X and Y scanning is shown (ie, directions are arranged transversely at, for example, 90 degrees to each other). Fig. 10D Shown are the parameters used in the remelting of the layer using a double scanning strategy.

[0134] (2) Alternative scanning: Figures 11A-11D As shown, alternative scanning strategies can be implemented, for example, by using the Realiser SLM 100 system. Such alternative scanning strategies may include Fig.11A and 11B The X and Y alternating hatched laser gratings shown and Fig. 11C and 11D Circumferential laser grating shown.

[0135] (3) Online Monitoring: Ultra-low oxygen levels can be maintained in the build atmosphere. When processing reactive materials, oxygen concentrations below 50 ppm can be critical and can contribute significantly to material integrity and mechanical properties. For real-time melt pool monitoring systems, SLM and electron beam qualification are performed by creating a database that can include various information such as laser power, scanning strategy, hatching strategy, and other information that can describe the impact of process part parameters on the mechanical properties of the constructed part (e.g., a constructed intramedullary nail).

[0136] (H) Post-processing of RMT components: Post-processing of RMT components may include but is not limited to the following steps or processes.

[0137] (I) Heat treatment: Heat treatment of parts built via an AM process can involve any combination of HIPPING (Hot Isostatic Pressing), stress relieving, and annealing steps.

[0138] (a) HIPPING: Hot isostatic pressing (HIP) can be used to reduce the porosity of the metal and improve the mechanical properties and machinability of the material. HIPPING can include the following steps:

[0139] (1) Evacuate / clean (e.g., 3 times to below 15 mb);

[0140] (2) Holding temperature: Typical HIP temperatures may be, for example, between about 920°C and about 1000°C, and the optimum temperature may be about 980°C + / - 10°C. If the HIP temperature is above 1000°C, the additively manufactured parts may be contaminated with nickel because the HIPPING equipment is typically made of nickel-based alloys. This is more pronounced when the HIP temperature approaches 1050°C. This can be achieved by (a) wrapping the parts in titanium (Ti) foil, (b) placing the parts on a recrystallized alumina plate or box so that the titanium (Ti) cannot come into contact with the nickel (Ni)-based load plate, or (c) placing the wrapped parts on a sawdust blanket;

[0141] (3) Maintaining pressure (MPa): For example, the pressure during at least a portion of the HIP process is 103 MPa + / - 5 MPa;

[0142] (4) Holding time (minutes): for example, 120 minutes + 15 / -0 minutes duration of the HIP process;

[0143] (5) Cooling rate (°C / min): for example, less than 10°C / min; and

[0144] (6) Heating rate (°C / min): For example, less than 10°C / min.

[0145] (b) Stress relief process: Stress relief can be performed in a stress relief furnace under an argon atmosphere or in a vacuum furnace. The stress relief process may include, among other steps, one or more of the following steps:

[0146] (1) raising the temperature of the additive manufacturing process to an elevated temperature, such as a temperature of about 800° C., in a Centorr vacuum furnace within 60 minutes;

[0147] (2) maintaining the elevated temperature of the additively manufactured build part for a predetermined period of time, such as about 2 hours; and

[0148] (3) When the temperature of the additively manufactured part drops to a set cooling temperature (e.g., a temperature of about 400°C), the furnace heating power is turned off and the furnace door is opened. The maximum cooling rate for reducing the temperature of the high-temperature additively manufactured part to the set cooling temperature may be, but is not limited to, 55°C / min, while the cooling rate from the set cooling temperature to the base temperature (e.g., a temperature of about 100°C) may be slower, such as about 35°C / min.

[0149] (c) Annealing: The annealing process may include the following steps:

[0150] (1) heating the fabricated part in an argon inert atmosphere at an annealing temperature, for example, about 1000°C for 2 hours. This temperature exceeds the β-transus temperature of the Ti-64 alloy, 995°C; and

[0151] (2) Nitrogen quench to room temperature. Note that the 1000°C heat treatment may be used to take the α phase into solution and fully sinter the component but may only partially sinter to adjacent powders.

[0152] (2) Machining operations: The surface modification techniques for the external and internal geometries of additively manufactured parts outlined below can provide finished surfaces with improved mechanical properties and reduce the risk of bacterial contamination by eliminating surface negatives typically inherent on most machined surfaces. Removing sharp edges from additively manufactured parts can also facilitate smoother, less disruptive introduction into the human body, where tissues may otherwise be damaged or traumatized by sharp edges, etc.

[0153] (a) Machining Operations - External Geometry: The following optional surface finishing steps can be used to remove the alpha shell from the exterior surface of the additively manufactured build part, which is typical of three-dimensional (3D) printed Ti-64 parts. The surface finishing operations can also help to flatten the surface, improve part accuracy, and introduce a compressive layer into the first 0.2 millimeters (mm) of the surface.

[0154] (1) Removal of alpha shell by sandblasting: The alpha shell is removed, which can be about 30 microns deep, but is not necessarily uniform. The alpha shell can be removed in a variety of different ways, including mechanical abrasion methods using alumina media. This step can be performed manually based on experience to observe the sparks generated by the alpha case to help determine when the substrate is destroyed (i.e., the spark will be extinguished when the alpha guest has been removed). In a production environment, automated groups can be established, which can in turn give more uniform metal removal. Mechanical removal has the advantage of helping to prepare the surface for subsequent operations, as well as being a lower cost option compared to chemical milling. Maintaining the geometry of the AM build part can also provide some challenges, as the material removal rate between the alpha shell and the substrate will vary significantly. This method can be used to attack both the outer and inner surfaces of the intramedullary nail. The material loss from this step (typically 0.2 microns (μm)) will be factored into the outer diameter (OD) and inner diameter (ID) of the associated model of the AM build part and / or CAD file.

[0155] (2) Vibratory polishing: Depending on the surface condition after shot peening, it may be necessary to rough grind the part to remove or truncate the peaks before shot peening. This step ensures that the shot peening process compresses the entire surface without the risk of folding over surface asperities that would otherwise create stress concentrations.

[0156] (3) Creating a compressive layer of residual stress: After removing the α shell and preparing the surface, the shot peening parameters are specified to induce an optimal compressive residual stress layer with a maximum value of about 800-1000 megapascals (MPa) and a depth of about 0.2 millimeters (mm). Fig.12 It is shown that the compressive layer will have a fine grain structure and will effectively delay the initiation and propagation of fatigue cracks. Specifically, Fig.12 A typical hammer head stress profile for a Ti-64 ALM component is shown. Shot peening of the sample surface has the effect of increasing the surface hardness, as well as introducing beneficial compressive residual stresses that reduce the tensile stresses felt at the surface.

[0157] (b) Machining Operations - Internal Geometry: The following optional surface finishing steps can be used to complete the internal surfaces of the part.

[0158] (1) Extrusion honing: Extrusion honing is an internal surface finishing process characterized by flowing an abrasive-containing fluid through the workpiece, which effectively performs erosion. This fluid is usually very viscous and has the consistency of putty or dough. It can be particularly useful for removing burrs, polishing surfaces, forming radii, and even removing material. The properties of AFM make it ideal for the internal surfaces of intramedullary nails, slots, holes, cavities, and other areas that may be difficult to reach with other polishing or grinding processes.

[0159] In another form of the invention, the fabrication of an orthopedic implant may utilize optimal processing conditions during component formation. In one embodiment, the fabrication of an orthopedic implant includes optimizing fatigue performance of a three-dimensional (3D) printed Ti-64 component.

[0160] (A) Laser Power: The mechanical properties of additively manufactured parts can depend on how much power is used to build them, i.e., the energy density of the laser beam used to produce the part. In general, the greater the energy density used to make a part, the rougher the surface finish of the part. This phenomenon can be caused by heat from the part "leaking" into the surrounding powder material and causing the powder to fuse to the surface of the part. Therefore, increasing the energy density of the laser beam can increase the surface roughness and overall strength of the part.

[0161] For example, a batch of twelve (12) additively manufactured Ti-64 samples similar to the distal segment of a tibial nail were formed by laser sintering. The following processing conditions were implemented using a Renishaw 250 ALM: (a) a scan speed of 150 millimeters per second (mm / s); (b) a focus offset of 0 millimeters (mm); (c) a spot distance of 65 micrometers (μm); (d) an exposure time of 250 μs; and (e) a laser power varied between 120 watts (W) and 400 watts (W). This processing includes and provides relevant information about:

[0162] (1) Metallography: Twelve ALM samples were exposed to varying laser powers of 120 W, 160 W, 200 W, 240 W, 280 W, and 350 W and then axially sectioned into four sections (S1-S4 ( Fig.13 )); distal to proximal), and according to Fig.13 The schematic diagram shown in the following metallurgical studies was used. Specifically, Fig.13 A three-dimensional (3D) model of an ALM test component attempting to clone the distal end of a standard intramedullary nail is shown.

[0163] After cutting, the sliced ​​Ti-6Al-4V ALM samples were cold mounted in acrylic resin and polished to 1 micron with SiC paper (80, 220, 800, 1200, 2400 grit). The porosity in the wall sections of the samples was measured from scanning electron microscopy (SEM) images using standard image analysis software. The porosity of each sliced ​​sample was determined in the wall section based on the SEM images captured using image analysis software. The results of the study are summarized in Fig.14 middle, Fig.14 shows the ALM samples (S1 to S4 ( Fig.13)). The porosity is between 0.03% and 1.7% in the wall section and tends to be highest towards the proximal end of the sample (S4). In general, the porosity decreases with increasing laser power, implying an improved consolidation of the powder structure.

[0164] SEM images captured from the fracture surface of a sintered ALM part laser driven at 400 watts (W) Fig.15 This is shown in Figure 1, which is characteristic of both ductile and brittle failure. This in turn indicates that there is some porosity in the core structure combined with some unsintered powder. Specifically, Fig.15 Shown is an SEM image captured of a fracture surface from an ALM part sintered to 400 Watt (W) laser. The morphology is mainly ductile, but the deformation around the pre-existing crack is rather poor due to the relatively poor machinability of the alloy.

[0165] (2) Four-point bending test: For the four-point bending fatigue test method of intramedullary fixation devices, refer to the ASTM test method (F1264-03). Samples exposed to varying laser powers of 140W, 180W, 220W, 260W, 300W and 400W were subjected to four-point bending cyclic fatigue tests. Two load conditions were used, namely 200 Newtons (N) to 2000 Newtons (N) and 300 Newtons (N) to 3000 Newtons (N). In addition, the processed Ti-64 nails were cut to a length similar to the shortest sample nail (77.04mm) and tested using the 300 Newtons (N) to 3000 Newtons (N) method to verify the test bench using relatively short nails. The processed Ti-64 nails survived 10 6 In all cases, the length / diameter (L:D) ratio was fixed at 7:1 and all tests were performed at 5 Hertz (Hz). The test results are summarized in Fig.14 and Fig.15 middle, Fig.16 shows the summarized test results performed at 5 Hz, Fig.17 The effect of laser power on the four-point bending fatigue performance of a Ti-64ALM component under a step load of 2000:200 Newtons (N) at 5 Hertz (Hz) is shown.

[0166] (B) Standard fatigue properties of laser sintered and electron beam melted specimens: To determine if there is any cumulative benefit when using a specific machine supplier to build Ti-64 nails, such as Fig.18As shown, it is possible to create part geometries that are relevant to the final device, which in turn allows for direct comparisons. Fig.18 List of machine suppliers for Laser Sintering and Electron Beam Melting of Ti-64 material, Fig.19A A 2D diagram of the simplified test specimen geometry is shown. Fig.19B Photoetching of a simplified test coupon geometry is shown.

[0167] When the ALM Ti-64 specimens outlined in Figure 19 were subjected to four-point bending fatigue testing in the “as-built condition” by different machine suppliers, as Fig. 20 As shown in Figure 1, they tend to fail at about 125k cycles. In contrast, wrought metal of the same geometry can run for 1M cycles when loaded at 4000 Newtons (N) to 400 Newtons (N) at 5 Hertz (Hz). When these ALM Ti-64 specimens were observed under a microscope, they were found to be highly porous, exhibit a rough surface, and have a suboptimal microstructure. More specifically, Fig. 20 Four-point bending fatigue performance of laser sintered and electron beam melted Ti-64 specimens as a function of machine supplier is shown. Part geometry: - 100 millimeters (mm) long x 10 millimeters (mm) OD and 4.7 millimeters (mm) ID); Test conditions: -2000 Newtons (N) to 200 Newtons (N) at 5 Hertz (Hz) according to ASTM 1264. This testing shows that the average performance of the laser sintered and electron beam melted specimens is about 125k cycles to failure.

[0168] refer to Fig.21A and 21B , which shows the typical crystal structure of a laser sintered Ti-64 part, which includes a band β within a matrix α and is of the Widmanstatten type. Specifically, Fig.21A and 21B Microsections of EOS samples at different magnifications are shown and etched with Kroll's reagent. This closely resembles a forged microstructure produced after heating to 1000°C and forced air cooling. The grain sizes are in the range of 100 micrometers (μm). This type of structure may not produce good mechanical properties and may require appropriate solution and aging heat treatments to produce a material with acceptable mechanical properties. EOS Ti-64 section etched with Kroll's reagent at low magnification ( Fig.21A ) and higher magnification ( Fig.21B ) shows the crystal structure, showing the Widmanstatten-type structure of the individual grains.

[0169] (C) HIPPING: Honing can very effectively improve the fatigue properties of ALM Ti-64 parts, such as Fig. 22 As shown. In general, when the ALM Ti-64 parts are subjected to HIPPED at 980°C for 4 hours at a pressure of 200 MPa with an initial cooling rate of 4°C and 10°C / min, the fatigue performance can be improved by 100%. For the samples produced by the SLM solution, when loaded between 4000 Newtons (N) and 400 Newtons (N) at 5 Hertz (Hz) (not shown), the ultimate failure occurred at 26,891 cycles. Specifically, Fig. 22 Four-point bending fatigue performance as a function of machine supplier is shown for HIPPED laser sintered and electron beam melted specimens with the following parameters: part geometry 100 millimeters (mm) long, 10 millimeters (mm) outside diameter, 4.7 millimeters (mm) inside diameter; test conditions 3000 Newtons (N) to 300 Newtons (N) at 5 Hertz (Hz) per ASTM 1264; HIPPED at a temperature of 980°C at a pressure of 200 MPa with an initial cooling rate of 10°C / minute for four (4) hours. Failure typically occurs at approximately 200,000 cycles, with the HIPPED part produced by the implementer being exhausted by 1M cycles, Fig. 22 .

[0170] The HIP process can refine the crystal structure to produce a lamellae structure that has some similarities to conventionally manufactured titanium when heated above the beta transus temperature and slowly cooled. Fig.23A and 23B Figure 1 shows the effect of HIP processing on grade 23 titanium produced on an EOS machine. Fig.23A shows the components of the construction, Fig. 23B The part after HIP treatment is shown. While the width of the lamellae and overall grain size vary with other samples, the illustration is typical of the type of modification observed on samples from other machines. The main difference with conventional Ti-6Al-4V titanium is the rather disordered nature of the lamellae observed in all of these samples. The microstructure observed with the laser sintered parts consists of bands of alpha separated by very small areas of beta.

[0171] like Fig.24 As shown, conventional lamellar Ti-6Al-4V titanium forms larger areas of parallel laminae within the crystal. Specifically, Fig.24 The typical ordered structure of conventional thin-plate Ti-6Al-4V titanium alloy is shown. Fig. 23B )Compare Fig.24The conventional lamellar type Ti-6Al-4V titanium alloy shown is more disordered. This contributes to the lower fatigue performance of these samples during fatigue testing. It can also be seen that there are well-defined crystal boundaries in the conventional samples, but these are difficult to define in the examined test samples as there does not appear to be a clear distinction between lamellar colonies and grain boundaries. More specifically, Fig.24 A Ti-6Al-4V titanium alloy is shown which is typically forged and heated to a beta transus temperature of 100°C, cooled at a rate of 100°C min-1 to 700°C, stabilized for 2 hours, and then cooled in an oven.

[0172] The ALM samples were HIPed at 980 °C and therefore most likely did not reach the β-transformation and therefore no phase transformation to β was achieved. This occurrence combined with a slow cooling rate of 10 °C min-1 explains the differences between the as-forged microstructure and the ALM samples studied. In order to produce a suitable structure by heat treatment, the β-transformation temperature must be reached. This will allow the creation of an ordered lamellae structure upon cooling. The width of the lamellae is determined by the cooling rate, with a finer structure being achieved at a faster cooling rate. It has been found that the width of the lamellae (t) should be 3 microns and the size of the colonies of parallel lamellae (d) should be 30 microns to obtain maximum fatigue strength.

[0173] If the HIP temperature is increased to 1000°C, a more ordered parallel structure should be achieved so that a complete beta transformation is achieved. In order to achieve an ordered structure of an α+β lamellae type alloy, the cooling rate from the beta transformation temperature should be between 0.24°C min-1 and 72°C min-1, with faster cooling rates producing smaller flakes. The cooling rate in the HIP furnace used for this batch of ALM parts was 10°C min-1, and the flakes produced had a width between 4.05 microns and 6.12 microns, Fig.26 This is not far from the optimum size for maximum fatigue strength. Therefore, if the HIP temperature is increased to 1000°C, thereby achieving a complete β transformation, a more ordered parallel structure should be achieved.

[0174] (D) HIPPING & MACHINING / POLISHING: Post-processing of the outer surface of a part has a significant improvement in fatigue performance. For example, techniques such as abrasive fluid machining result in surface flattening and reduce the number of crack initiation sites in the test part. Fig.25 As shown, the surface finish or roughness of the outer geometry of the manufactured part subjected to such post-machining increased from 5.43 μm (Realizer) - 23.4 μm Ra (Arcam) to 0.4 μm Ra after abrasive fluid machining.

[0175] The combined effects of HIPPING and external polishing produce components with fatigue properties comparable to those of forged components (i.e., exhausted at 1M cycles when loaded between 4000 Newtons (N) and 400 Newtons (N) at 5 Hertz (Hz), as shown in Figure 2. Fig.28 The results also show that ALM parts made from TiAl6V4 Grade 5 material (higher oxygen content) perform well. This result can provide significant benefits because Grade 5 powder material typically costs 35% less than Grade 23 material, making the economic justification of this method relatively significantly more attractive. Specifically, Fig.28 Shown are the surface finish and four-point bending fatigue properties of HIPPED laser sintered and electron beam melted specimens as a function of machine vendor with the following parameters: part geometry with 100 millimeters (mm) length, 10 mm outer diameter, and 4.7 mm inner diameter; and test conditions from 4000 Newtons (N) to 400 Newtons (N) at 5 Hertz (Hz) according to ASTM 1264. Figure 29-30 Load extension curves for heat treated and non-heat treated ALM parts are shown. Points F and J represent the performance of the non-heat treated parts. These parts are strong but brittle.

[0176] The present invention provides at least the following advantages over the prior art: (1) mechanical properties similar to forged components; (2) grade 5 Ti powder can be used, which in turn can provide economic advantages; and (3) improved part tolerances. However, it should be understood that these advantages are exemplary and do not limit the scope of the present invention in any way.

[0177] Embodiments of the present invention also provide an optimal DMLS manufacturing route for orthopedic devices, which may be manufactured in 1 / 6 the time. In another embodiment, a method for manufacturing an elongated orthopedic device by direct laser sintering is provided, comprising the following steps:

[0178] (1) Manufacturing a virtual 3D external slender device model;

[0179] (2) using a laser power of at least 300 watts (W) and TiAl6V4 powder of at least grade 5 quality, to fabricate an elongated device in a suitable build orientation according to the 3D model by direct metal sintering, and wherein the powder is sintered to a set diameter only around the outer and inner circumferences of the model, such that a central portion between the inner and outer circumferences consists essentially of unsintered powder;

[0180] (3) Scanning at a speed greater than 4 times the normal speed of laser sintering (i.e., 3000 mm / s);

[0181] (4) hot isostatic pressing (HIP) the elongated ALM device using a temperature of at least 1000°C with a cooling rate between 0.24°C min-1 and 72°C min-1, thereby ensuring sintering of the central powder segment; and

[0182] (5) The HIP-processed elongated device is machined and polished to the desired geometry and surface tolerance (e.g., 32 Ra μm (Realizer)).

[0183] The following two scanning strategies ("Ultra" Laser Scanning and "Boundary Scan - In-Situ Bombardment") assume that most of the cost of goods for manufacturing ALM nails is associated with the build time in the DMLS machine.

[0184] (A) “Ultra” laser scanning: The economic advantages of operating a laser sintering machine at “ultra-scanning speeds” are Fig.31 and 32 This is evident from the data depicted in Figure 1, which was generated by three DMLS machines equipped with single, dual, and quad scanner optics. For a conventional laser scanning program for a single laser system, a batch of 100 nails would take approximately 4.5 days to complete, which is equivalent to 65 minutes per nail ( Fig.25 This does not compare to the throughput obtained from a standard machining operation, which is relatively equivalent to about 2.5 minutes per nail or 600 nails per day. If a DMLS machine equipped with four lasers is set to operate at a higher scan speed, each nail would only take 12 minutes to build ( Fig.32 ), thereby making the process significantly more attractive for high-volume manufacturing (i.e., 10,000 parts per year). By operating at higher scan speeds (typically 3-4× standard speeds for laser sintering; 1000mm / s), the part will only be partially sintered (typically 50% of the entire volume of the part). This accelerated scan speed is still significantly lower than the scan speeds used in the electron beam melting process (1000-8000 m / s). This scanning strategy assumes the following: (a) any residual porosity left after the build is removed after heat treatment; (b) a standard heat treatment cycle is used; and (c) dimensional changes due to part shrinkage are accounted for in the original design of the CAD file or by Magic Laser. ® program control; and (d) the mechanical properties of the overscan speed portion are equivalent to those of the normal scan.

[0185] Fig.31 Calculations are shown for generating 100 fully dense intramedullary nails using single, dual, and quad lasers at normal scan speed and with the following parameters: slice count = 3911; recoat time = 31,288 seconds; scan speed = 600 mm / s. Fig.32Calculations for producing 100 fully dense intramedullary nails at "hyperscan speed" are shown, and utilize the following parameters: slice count = 3911; recoat time = 31,288 seconds; and scan speed = 3000 mm / s.

[0186] Figures 33A-34B Figure 1 shows the breakdown of the cost for manufacturing an additive manufactured Ti-64 nail. Fixed costs include the cost of powder (assuming current supply chain selection of qualified powders, without access to lower cost metal spray equipment), post-processing, and packaging / sterilization. The total cost of manufacturing the ALM nail using this manufacturing route is $149.50. The cost of operating at superscan speeds has a significant impact on the build time cost, reducing it from 62% of the total manufacturing cost to 25.6%, as shown in Figure 1. Fig.33B and Fig.34B In addition, the cost of producing the “superscanned” part was reduced from $149.50 to $75.90, which is consistent with the cost of making a machined part of equivalent geometry.

[0187] The production cost of manufacturing 100 nails using standard scanning conditions of SLM DMLS machine (scanning speed <1000mm / s) is Fig.33A is shown as a cost breakdown ($ / manufacturing step), and Fig.33B As a percentage cost for each manufacturing step, with the following assumptions: the cost of running the ALM is $97 / hour; the cost of implant grade powder is $255 / kg; and the build time is 96 hours to make 100 nails.

[0188] Fig.34A and Fig.34B The production cost of manufacturing 100 nails using the “higher speed” scanning conditions (scanning speed < 1000mm / s) for an SLM DMLS machine is shown, showing the cost breakdown for each manufacturing step ( Fig.34A ) and the percentage cost of each manufacturing step ( Fig.34B ), and using the following assumptions: the cost of running the ALM is $97 / hour; the implant-grade powder cost is $255 / kg; the build time is 96 hours to make 100 nails; and the heat treatment cycle is HIP at 1050°C for 6 hours, followed by cooling in air as quickly as possible, followed by a heat treatment anneal at 840°C for 6 hours, where all heat treatments are performed in an inert argon atmosphere.

[0189] (B) Boundary Scanning: A second approach for reducing the cost of building parts in a DMLS chamber is to use a scanning strategy that limits sintering to the boundary layers (i.e., the outer and inner surfaces of the cannula portion), aptly referred to as a "baked bean can or in-situ shelling" model. This scanning strategy omits the cross-hatching or core scanning of the part. A 3D model of the proximal end of the TrigenMeta tibial nail highlighting the concept of boundary layer scanning is shown in Fig.35 Using this scanning method, only 5% of the total volume of the part is sintered on the build platform. Specifically, Fig.35 Concept showing a boundary or "baked bean can" scanning strategy used to densify the outer and inner layers of a TriGen Met tibial nail (proximal tapered region) to reduce build time in the DMLS chamber. The core of the part contains a free-flowing powder that is densified after heat treatment. In this example, certain part design features, such as keyways and slotted screw holes, have been closed.

[0190] Additionally, in certain embodiments of the present invention, an intramedullary nail or other type of implant is manufactured by the following steps and processes.

[0191] (I) Boundary Scan: Figure 36A and Fig.36B A cross section of an intramedullary nail 110 including an additively manufactured component 100 after boundary scanning is shown, highlighting the resized portion to account for shrinkage effects ( FIG. 36A ), a HIP processed additively manufactured component indicating the final “stable” nail 110 dimensions after heat treatment ( Fig.36B ). Specifically, FIG. 36A shows an additively manufactured component 100 that will provide an orthopedic transmedullary nail 110 prior to boundary scanning, indicating a CAD file that is sized to account for dimensional shrinkage (typically <5%), and Fig.36B An orthopedic intramedullary nail 110 after HIPPING is shown, indicating a fully dense part with a specified geometry. Referring to FIG. 36A , an additively manufactured component 100 is manufactured using DMLS technology with a boundary-only scanning method. In conventional practice, various scanning strategies are used to produce a sintered model representing the complete intramedullary nail device 110. In the proposed method, a "baked bean can" of the additively manufactured component 100 can be manufactured by laser sintering the three-dimensional model, more specifically the outer surface 102 and the inner surface 104 to a desired thickness, which can be, for example, in the range of about 100 micrometers (μm). In addition to the unsintered pressed powder core 108, the internal cannula 106 will remain using this scanning method. Reference Fig.36B After HIPPING, the nail 110 becomes dense by heat treatment with some accompanying dimensional changes, which reflects the amount of shrinkage experienced by the additively manufactured part 100 during the heat treatment step. Typically, using e.g. Fig.37AThe boundary scan strategy shown provides the powder core for the nail. The part was densified in the build chamber at 5%. The bottom surface of the nail was extruded to 3mm in Magics software to prevent powder from escaping from the far end of the part after removal from the platform. Fig.37B . Fig.38 The porosity remaining in the cross section of the component is shown. Typically, the porosity measured is 0.3%, Fig.38 In addition, the average pore size is about 3.4 microns, Fig.38 .

[0192] (II) Super laser scanning: Figure 39A and Fig.39B A cross section of an intramedullary nail 206 comprising an additively manufactured component 200 is shown after an ultra-laser scan has been produced to produce 50% densification in the part. The figure highlights the size adjustment of the nail 206, which accounts for shrinkage effects on the additive manufactured component 200, after heat treatment (FIG. 39A), which may occur, and the HIP treated nail 210 indicating the final "stabilized" part size after heat treatment ( Fig.39B ). Specifically, FIG. 39A shows an additively manufactured component 200 to be used to form an orthopedic transmedullary nail 210 after super laser scanning indicating a sizing CAD file to account for dimensional shrinkage, and Fig.39B The orthopedic posterior intramedullary nail 210 is shown after HIPPING, showing a fully dense component having the desired geometry.

[0193] Referring to FIG39A , an additively manufactured component 200 is manufactured using DMLS technology with ultra-laser scanning. In conventional practice, various scanning strategies are used to produce a sintered model representing a complete intramedullary nail device 210. In the proposed method, the cannula 202 is surrounded by a semi-sintered core 204. See Fig.39B , after HIPPING, the nail 206 is densified by heat treatment with some accompanying dimensional changes that reflect the amount of shrinkage experienced by the part during the heat treatment step.

[0194] The super-scanned parts provided by the SLM solution were subjected only to heat treatment and the number of cycles to failure was compared with the standard-scanned parts produced after the same steps in the process, Fig.40The overscanned parts consisted of three groups of parts: full skins with pressed powder cores, full skins with pressed powder cores with extruded distal ends to allow the parts to be heat treated on the build platform, and parts with full skins and partial powder cores. When loaded between 3000 and 300N, the average cycles to failure were 1,161 + / - 288 for the parts with full skins and pressed powder cores after the heat treatment step. The corresponding average cycles to failure were 34.7 + / - 9.6k for the parts with full skins and pressed powder cores with extruded distal ends after the heat treatment step. The corresponding average cycles to failure after the heat treatment step was 180,068 + / - 38.7k for the parts with full skins and partial overscanned cores with extruded distal ends. The improvements observed in the fatigue performance of the pressed powder core samples after the inclusion of the extruded distal ends are believed to be a benefit resulting from heat treatment away from the component plate. The extruded “bottom plate” prevents the powder from coming out. The pegs were positioned above the platform and then dropped 2mm (in CAD space). It was assumed that HIP would not work well with the pegs without the extrusion floor as there would be stresses and micro cracks around the base which would make them rather "leaky". For comparison purposes, parts subjected to a standard scanning strategy (EOS, Arcam and SLM Solutions) are also included in this figure. The data shows that the fatigue performance of the partially sintered cores is comparable to that of the EOS supplied parts subjected to a standard scanning strategy, Fig.40 .

[0195] In another embodiment of the present invention, an antimicrobial staple may be provided.The cost savings realized from the scanning strategies discussed above may enable other cost-effective processes to be included within the process map.

[0196] In another embodiment of the present invention, a silver hip implant may be provided. Specifically, Fig.41 As shown, during the HIPPING step, silver is deposited on the hip implant and the following methods can be used: (a) silver plated / coated workpiece; and (b) modified compressed medium (argon-silver vapor). Both methods use non-line-of-sight methods to produce silver-coated products, which are known to have antimicrobial properties. Specifically, Fig.41 A sagittal section through a HIPPING furnace is shown, highlighting the exposure of the ALM portion to the silver vapor. This coating technique assumes that the deposited silver layer is thicker than 0.1 millimeter (mm) to accommodate any post-processing operations.

[0197] In another embodiment of the present invention, a silver coated bead shot peening process may be provided. Fig.42 As shown, this embodiment can constitute a one-step process involving the use of silver-coated ceramic particles prepared by firing a silver plating or silver nitrate coating on the surface. In addition to producing the desired surface finish, the particles produce a hardened layer of silver-titanium dioxide with antimicrobial properties. Specifically, Fig.42 is a schematic diagram of the bead jetting process highlighting the deposition of silver on a work part during a finishing operation using a silver coated mesh or bead jetting.

[0198] The present invention provides significant savings in terms of cost of goods relative to existing implants and manufacturing methods. However, it should be understood that these advantages are exemplary and do not limit the scope of the present invention in any way.

[0199] Another embodiment of the present invention provides an orthopedic implant device, such as an intramedullary nail having the design freedom of additive manufacturing, and which includes a longitudinal internal channel capable of accommodating a removable sensor probe, wherein the removable sensor probe is configured to align with a distal locking hole and a proximal locking hole. The orthopedic implant device may also have an internal geometry that promotes variable stiffness in the anterior-posterior (A / P) plane and the medial-lateral (M / L) plane and / or an internal geometry that provides a lower stiffness implant for larger patients. Another embodiment of the present invention provides a minimally invasive method for automated mechanization to provide biomechanical loading to a healing fracture.

[0200] (A) In-situ distal and proximal sensor probes for intraoperative screw hole targeting: Embodiments of the present invention include longitudinal internal channels created in a wall portion of an orthopedic device (e.g., an intramedullary nail) using additive manufacturing to accommodate removable sensor probes for registering distal and proximal screw holes. For example, Fig.43A An intramedullary nail 300 is shown having an internal sensor probe channel 302 extending parallel to a cannula 304. In addition, Fig.43B A microCT image captured from an ALM part with built-in channels in the wall section is shown. A longitudinal section through the part shows that there is no residual powder in the channels after the build stage. Fig.48 The proximal end of an intramedullary nail 300 constructed by additive manufacturing is shown and includes an internal sensor probe channel 302 ( Fig.43A ), which is adapted to receive the insertion of a sensor probe 308. Including a removable sensor probe in channel 302 can facilitate proximal locking because the sensor probe is not located in the cannula of the nail, thereby allowing the surgeon to drill the nail and insert the screw into the nail 300.

[0201] The shape of the internal sensor probe channel 302 can be created using additive manufacturing based on the constraints imposed by the geometry of the intramedullary nail 300. In the illustrated embodiment, the internal sensor probe channel 302 is located within the wall portion 306 of the intramedullary nail 300 to ensure that a probe that can be accommodated in the channel 302 will not be trapped in the intramedullary cavity during removal. The longitudinal channel has a diameter of approximately 1.5 mm, extends the length of the nail and terminates just above the upper distal screw hole. Positioning the internal sensor probe channel 302 within the wall portion 306 will avoid the need for a welded cover, which would otherwise increase the complexity and cost of the manufacturing process.

[0202] Fig.44 An embodiment of a removable sensor probe designed to operate in this channel 302 is shown. Typical probe and sensor dimensions may include, but are not limited to, 1.422 millimeters (mm) outer diameter, 1.22 millimeters (mm) inner diameter, 1.12 millimeters (mm) height, 0.61 millimeters (mm) width, and 30 millimeters (mm) length. In addition, the sensor probe may be constructed of a variety of different materials, including, for example, stainless steel or another rigid or semi-rigid metal material. It may also include a potting compound, such as medical grade silicone rubber or epoxy, to protect the electronic components from moisture and vibration forces. In addition, the configuration of the sensor components on the printed circuit board (PCB), i.e., two wound ferrites (also known as six-degree-of-freedom tracking sensors), are preferably arranged 180 degrees from each other. Such a configuration may minimize the overall diameter of the sensor unit, for example, to 1 millimeter (mm), and it may be located in the internal sensor probe channel 302, which is less than 1.5 mm in diameter in the nail wall. The location of the channel also allows the nail to be locked at the distal end or the proximal end, first providing more options for the surgeon. The probe is also designed to be removed after the nail has been locked within the bone channel 302. Fig.45A and Fig.45B A model of an intramedullary nail 400 is shown having a channel 402 formed in an outer surface 404 of the nail 400 at the proximal end suitable for conventional manufacturing techniques. As depicted, Fig.45B The channel in includes a removable sensor probe 406. This design requires a welded plate to prevent the probe from becoming blocked during extraction from the bone canal.

[0203] Fig.46A and 46B An exemplary geometry of an internal sensor probe channel 302 that can be produced using additive manufacturing is shown, and the location of the channel 302 relative to the neutral axis is highlighted. According to certain embodiments, the size of the internal sensor probe channel 302 can be minimized to mitigate the risk of adverse fatigue failure. Including the channel 302 can reduce the bending stiffness in the plane of the channel 302 by approximately 11%, which is unlikely to cause any problems clinically. In addition, in the model shown, as shown in FIG. Fig.46AAs shown, the "x" axis in the plane of the channel 302 experiences an approximately 11% reduction in moment of inertia. Additionally, the "y" axis perpendicular to the channel 302 can experience an approximately 1% reduction in moment of inertia. Furthermore, the proximal end of the internal sensor probe channel 302 is modeled, and the moment of inertia with and without the channel 302 is calculated to be X: 878.8 mm 4 ; Y: 990.5mm 4 And X: 993.9mm 4 ; Y: 993.9mm 4 .

[0204] Fig.43A The intramedullary nail 300 shown in FIG. 1 can be manufactured in a variety of different ways, including, for example, by using laser or electron beam three-dimensional (3D) printing. Examples of laser scanning conditions for a Realiser SLM100 system are shown in FIG. Fig.47 The thickness of the wall 306 of the intramedullary nail 300 can be increased, for example, to increase the wall 306 in order to at least attempt to prevent the internal sensor probe channel 302 from deforming during heat treatment (HIPPING). For example, the wall 306 of the intramedullary nail can be increased so that the outer diameter of the intramedullary nail is increased by 0.5 millimeters (mm). According to such an embodiment, the additional thickness of the wall 306 can be sacrificed during post-processing operations.

[0205] Additionally, the use of additive manufacturing may allow the intramedullary nail 300 to be manufactured without the use of exit points or openings that may be associated with the removal of unsintered powder from the internal sensor probe channel 302. More specifically, microCT images acquired from the test intramedullary nail 300 after three-dimensional (3D) printing indicated that the internal sensor probe channel 302 was not contaminated by residual unsintered powder. Fig.43B Thus, according to certain designs, the intramedullary nail 300 may not include such a removal point or outlet in fluid communication with the internal sensor probe channel 302. The absence of a removal point or outlet, which may at least be suitable for removing residual non-sintered powder, may help simplify the design of the internal sensor probe channel 302 and the manufacture of the intramedullary nail 300.

[0206] (B) Automated Nail: One of the fundamental concepts in orthopedics is the understanding that appropriate mechanical loading accelerates fracture healing. This is based on a process of adaptation, according to which the skeleton structure is constantly optimized in response to the mechanical environment, and it occurs in response to dynamic rather than static loads. More specifically, it is related to peak strain amplitude and load frequency. Although conventional intramedullary nails allow weight bearing forces to be applied thereto, they often isolate the fracture forces from the compression forces due to the presence of a locking screw, the main purpose of which is to prevent rotation. In addition, due to the rigid structure of the intramedullary nail, the intramedullary nail can actually lead to non-healing fracture situations as a result of the fixed distance between the fracture ends and the constant load share throughout the healing cycle applied to the fracture.

[0207] Axial dynamization of conventional statically locked intramedullary nails involves removal of one or more interlocking screws two to three months after initial surgery in an outpatient setting. This method requires an invasive procedure and typically has a resolution of approximately 1 to 5 millimeters (mm), which is typically determined by the width of the slot in the intramedullary nail and may be available in only one section of the intramedullary nail. An auto-dynamic intramedullary nail would overcome some of these deficiencies and provide a step-wise improvement in accelerated bone healing by continuously adjusting the load sharing applied to the fracture site. In addition, an auto-dynamic intramedullary nail may help prevent the occurrence of delayed union or nonunion of a fracture by allowing appropriate axial movement of the broken ends of the bone toward each other.

[0208] According to certain embodiments, developing an automatically functionalized intramedullary nail may include the following:

[0209] (1) Retractable intramedullary nail assembly: Fig.49 A perspective view of the distal end of a power intramedullary nail 500 according to an illustrated embodiment of the present invention is shown. As shown, the power intramedullary nail 500 includes a first portion 502a and a second portion 502b. In addition, the power intramedullary nail 500 is configured such that the axial position of at least one of the first portion 502a and the second portion 502b can be adjusted relative to at least another of the first portion 502a and the second portion 502b. This relative change in the axial position of the first portion 502a and / or the second portion 502b can be adapted to at least adjust the total length of the power intramedullary nail 500.

[0210] exist Fig.49 In the illustrated embodiment, the dynamic intramedullary nail 500 includes a telescopic portion 504 adapted to facilitate relative axial displacement of a first portion 502a and / or a second portion 502b of the dynamic intramedullary nail 500 and to maintain the first and second portions 502a, 502b. Fig.53AAs shown, according to some embodiments, the telescoping portion 504 may be the other of the first and second portions 502a, 502b of the sleeve that extends from an end of the first portion 502a or the second portion 502b and is received in the interior region 506a, 506b of the first portion 502a or the second portion 502b. Fig.53A In the illustrated embodiment, the size of the end 503 of the outer wall 505 of the first portion 502a can be reduced such that a portion of the outer wall 505 is sized to be received within the interior region 506b of the second portion 502b. In addition, according to certain embodiments, the diameter of the outer wall 505 at the end 503 of the first portion 502a can be reduced by 1.5 millimeters (mm) in diameter such that the outer wall 505 at the end 503 has a diameter no greater than 15 millimeters (mm), which can be less than the diameter of at least a portion of the portion of the interior region 506b in which the end 503 will be received. Alternatively, as Fig.53B As shown, the telescoping portion 504' may be a separate component slidably received within the interior regions 506a, 506b of the first and second portions 502a, 502b.

[0211] In addition, if Figures 50A-50C As shown, the telescopic portion 504 can be located at various different positions along the dynamic intramedullary nail 500. For example, as shown in the figure, the telescopic portion 504 can be located at the distal region ( Fig.50A ), the middle part ( Fig.50B ) or the proximal region ( Fig.50C ). This variety of positioning can facilitate continuous dynamic loading of different types of fractures.

[0212] According to some embodiments, the distal end 508a and / or the proximal end 508b of the telescoping portion 504, 504' can include one or more guides or feet 510 that can help keep the first portion 502a or the second portion 502b along or, according to other embodiments, the guides or feet 510 extend from an inner wall 514 of the first or second portion 502a, 502b in which the telescoping portion 504 is slidably received. According to other embodiments, the outer wall 516 of the telescoping portion 504, 504' can be sized relative to the mating inner region 506a, 506b in which the telescoping portion 504 is slidably received so as to prevent the first and second portions 502a, 502b from moving along the central longitudinal axis 512.

[0213] refer to Figures 53A-55, wherein the interior region 506a, 506b that slidably accommodates the telescoping portion 504 may include one or more protrusions 518 that are configured to limit the axial displacement of the telescoping portion 504. For example, according to certain embodiments, the protrusion 518 may be positioned to at least limit the extent to which the first and / or second portions 502a, 502b may be axially displaced in a manner that reduces the length of the dynamic intramedullary nail 500. Furthermore, according to certain embodiments, the protrusion 518 may be positioned within the interior region 506a, 506b in a manner that limits the extent to which one or both of the first and second portions 502a, 502b may increase the length of the dynamic intramedullary nail 500. The protrusion 518 may take a variety of different forms, including, for example, washers that are held in place by mechanical fasteners 520 (e.g., screws or pins) and other protrusions 518 and fasteners 520. According to Fig.54 and Fig.55 The embodiment shown, as Fig.55 As shown, the protrusion 518 can be a pin extending through one of the first and second portions 502a, 502b and received in a slot 522 of the outer wall 516 of the telescoping portion 504, 504'. According to such an embodiment, the size of the slot 522, such as the length, can limit the distance that one or both of the first and second portions 502a, 502b can be displaced relative to each other. Fig.53A As shown, according to some embodiments, axial displacement may be limited by a shoulder 524 of the first or second portion 502a, 502b that is larger than the size of an adjacent interior region 506a, 506b of the other of the first or second portions 502a, 502b.

[0214] (2) Actuator-controlled actuation: The powered intramedullary nail 500 may be equipped with at least one mechanical actuator 526 that may bias and / or influence the orientation of the powered intramedullary nail 500. For example, according to some embodiments, the actuator 526 may be a spring that applies force to the opposing inner regions 506a, 506b of the first and second portions to extend or compress the length of the powered intramedullary nail 500. Fig.53A According to some embodiments, the actuator 526 can be a spring configured to exert a force against the shoulder 524 of the first portion 502a and against the end wall of the second portion 502b. Fig.53B 45, according to other embodiments, the actuator 526 can apply force to the end walls 528a, 528b of the first and second parts 502a, 502b, respectively.

[0215] Additionally, the actuator 526 may be a spring that may be linear or nonlinear in nature, such as Fig.51As shown. Nonlinear or variable rate springs can provide enhanced control and support to prevent excessive impact forces from being applied to the bone ends compared to linear springs. Examples of nonlinear springs include, but are not limited to, conical springs, and variable pitch or variable diameter wire springs. Equation 1 represents the general nonlinear spring force displacement relationship:

[0216] K = Ko + Kn (Formula 1)

[0217] Where K is the spring stiffness, Ko is the constant rate (linear portion), and Kn is a function of displacement (non-linear portion). Additionally, according to certain embodiments, the actuator 526 may further include a buffer tank, which may be used, for example, with the spring to provide a more uniform and stable displacement to stabilize bone end displacement during weight bearing.

[0218] Actuator 526 or a combination of actuators 526 (including springs in combination with other types of actuators) can promote or otherwise allow axial movement comparable to the displacement that can be caused by conventional dynamization of the distal screw. For example, according to certain embodiments, actuator 526 can promote an axial displacement of approximately 1 to 5 millimeters (mm). In addition, such a plate-mounted actuator 526 can also provide a controlled cyclic compression force, control the size of the gap between the ends of the fractured bone, and provide an adjustable intramedullary nail hardness, thereby at least helping to make the dynamic intramedullary nail 500 more suitable for the surrounding bone as healing proceeds.

[0219] Fig.52A and 52BA schematic diagram of a powered intramedullary nail 600 for actuating loads in long bone fractures is provided, highlighting the degradation of a resorbable polymer. The degradation of the resorbable polymer can also allow the embedded actuator 626 to simultaneously release one or more bioactive agents that can stimulate fracture healing and / or reduce infection at any time. According to certain embodiments, the actuator 626 can be a spring (constant or variable rate) contained in a resorbable biocompatible polymer encapsulation that simultaneously controls spring compression or a glassy non-resorbable polymer, such as when exposed to an externally applied induction heater. Examples of suitable resorbable biocompatible polymers can include, but are not limited to, poly(caprolactone), poly(lactic acid) and poly(glycolic acid), as well as other resorbable biocompatible polymers. The degradation of the polymer can place the actuator 626 in a relaxed state of the actuator 626, thereby providing a continuous cyclic load to the fractured bone 632a, 632b. Alternatively, relaxing the glassy polymer by induction heating will also allow the spring to respond under a cyclic force. According to this situation, when force is applied to the construct, the ends 630a, 630b of the fractured bones 632a, 632b can be supported. When the polymer package can no longer withstand the load as a result of induction heating or degradation, the bones 632a, 632b can move relative to each other. If the induction heater is removed from the limb, the glassy non-resorbable polymer can prevent the movement of the spring, allowing the polymer to crystallize and harden. The mechanism can be optimized to promote only compression of the fracture gap 634 so that bone healing is not hindered, and if the fracture gap 634 is stretched and the bone ends 630a, 630b are pulled apart, bone healing can be delayed.

[0220] Alternatively, the actuator 626 can be encapsulated in its extended state, allowing the two pieces of bone 632a, 632b to be drawn together as the polymer encapsulation degrades. The actuator 626 and / or the polymer encapsulation can also be filled with active agents or molecules to help promote fracture healing and / or reduce bacterial colonization of the implant using growth factors. Such active agents can include, but are not limited to, heavy metal ions, such as gold and silver.

[0221] According to another embodiment, the polymer encapsulated actuator 626 is periodically activated by external application of physical energy (e.g., heat, ultrasound, or electricity). Such activation can help change the physical properties of the polymer encapsulation, such as Young's modulus of bending modulus and other properties. This change in the physical properties of the polymer encapsulation can change the polymer encapsulation to a state in which the polymer encapsulation at least helps to prevent or otherwise resist the relative axial displacement of the first portion 602a and / or the second portion 602b of the dynamic intramedullary nail 600, and can additionally cause compression or expansion of the length of the dynamic intramedullary nail 600. Therefore, the movement of the actuator 626 can be provided as needed.

[0222] The telescopic portion 504, 604 of the powered intramedullary nail 500, 600 may be developed such that it provides unidirectional or bidirectional translation, for example, as Fig.53A and 53B In addition, the rotation of the telescoping portion 504, 604 and other components of the dynamic intramedullary nail 500 coupled to the telescoping portion 504, 604 can be prevented by the groove 530 in the inner and outer directions of the inner portion or sleeve 532. Examples of similar grooves 820 and matching recesses 822 are also shown. Fig.58A Alternatively, the insertion of the aforementioned protrusion 518 (e.g., a screw or pin) extending through the machined slot 522 in the inner portion of the dynamic intramedullary nail 500, 600 can be used to prevent rotation of the first and second portions 502a, 502b while keeping the mechanical actuator 526 located within the recessed area of ​​the dynamic intramedullary nail 500. In this case, the protrusion 518 can also be mechanically adjusted by the surgeon to control the displacement of the mechanical actuator 526. In addition, the protrusion 518 can be located in the dynamic intramedullary nail 500 to prevent the fracture gap from increasing.

[0223] Fig.56 The effect of wall thickness on mechanical stiffness measurements of three intramedullary nails is shown, namely (1) a standard nail, (2) a nail having a machined recess in a cross section of the outer diameter of the nail that provides a 25% and (3) a nail having a machined portion in a portion of the inner diameter of the nail that reduces the inner diameter of the nail in the recessed portion of the nail and increases the size of the inner diameter of the nail by 25%. More specifically, Fig.56 The stiffness measured for each of the three above-identified nails of four outer diameters, namely, 13, 11.5, 10, and 8.9 millimeters (mm) in diameter, is shown. The results indicate that machining an external recessed portion in the outer surface of the intramedullary nail to form a telescoping portion has a more significant effect on reducing the mechanical bending stiffness of the nail than increasing a portion of the inner diameter of the intramedullary nail. For example, for a 13 millimeter (mm) diameter nail, increasing the size of the inner diameter of a portion of the nail reduced the stiffness of the nail from 42,000 pounds per square inch (PSI) experienced by a standard nail to a stiffness of 29,000 PSI, as experienced by a nail that included a machined internal region with a larger inner diameter, which was further reduced to a stiffness of 7,000 PSI for the nail that included a recessed portion in the outer diameter. Thus, compared to an actuator positioned on the outer surface of the powered intramedullary nail 500, 600, increasing the size of the inner diameter of a portion of the nail reduced the stiffness of the nail from 42,000 pounds per square inch (PSI) experienced by a standard nail to a stiffness of 29,000 PSI, as experienced by a nail that included a machined internal region with a larger inner diameter, which was further reduced to a stiffness of 7,000 PSI. Thus, compared to an actuator positioned on the outer surface of the powered intramedullary nail 500, 600, increasing the size of the inner diameter of a portion of the nail reduced the stiffness of the nail from 42,000 pounds per square inch (PSI) experienced by a standard nail, to a stiffness of 29,000 PSI, as experienced by a nail that included a machined internal region with a larger inner diameter, which was further reduced to a stiffness of 7,000 PSI. Figures 52A-53B The actuator 526, 626) shown in FIG. 5 is housed in the interior of the powered intramedullary nail 500, 600, which has less impact on fatigue performance and can be constructed by using additive manufacturing.

[0224] (3) Actuation - Controlled actuation: According to certain embodiments, as described above with respect to at least Figures 52A-53B The actuator 526, 626 discussed may be a bushing configured to locally generate a cyclic load at the fracture site. The bushing may be constructed of a variety of materials, including, for example, elastomeric materials such as rubber, as well as other materials. Furthermore, according to certain embodiments, the bushing may replace at least some types of actuators 526, 626, such as variable springs. Furthermore, as Fig.57 As shown, the bushings 700a-c can have a wall 702 that generally defines an interior region 704 of the bushing and is sized to receive at least a portion of the first and / or second portions 502a, 502b, 502a, 602b. Fig.57 As shown, the exterior of wall 702 can have a variety of shapes and sizes.

[0225] (4) Actuation - Spring and Shape Memory Collar Assembly: According to certain embodiments, the dynamic intramedullary nail 800 may further include an activatable shape memory sleeve or collar 802, such as Figures 58A-58C As shown, it is configured to control the actuator 804, for example, to control the displacement experienced by the spring actuator and the resultant force applied to the bone. Figures 58A-58C As shown, according to some embodiments, a sleeve or collar 802 can be positioned around a telescoping portion 806 of a powered intramedullary nail 800, and the telescoping portion 806 is sized to be accommodated in an interior region 812 of a first portion 810a and / or a second portion. According to the illustrated embodiment, a first end 814a of the sleeve or collar 802 can be positioned adjacent to an adjacent end of the actuator 804, and a second end 814b of the sleeve or collar 802 can be positioned adjacent to an adjacent portion of the second portion 810b. In addition, in Figures 58A-58C In the illustrated embodiment, the sleeve or collar 802 and the actuator 804 are located in the distal region of the powered intramedullary nail 800. However, the sleeve or collar 802 together with the actuator 804 may be positioned at various other locations along the powered intramedullary nail 800.

[0226] The sleeve or collar 802 can be a shape memory ring, such as a shape memory polymer or metal alloy, which is stimulated to contract when activated, such as when heated above body temperature, to accommodate fixed translation. Alternatively, the sleeve or collar 802 can be constructed of a piezoelectric material that deforms in a manner that increases or decreases in size in response to an externally applied voltage when activated from an inactive state to an active state. Thus, the sleeve or collar 802 can provide a locking mechanism for the automatically actuated powered intramedullary nail 800 when the sleeve or collar 802 is adjusted to be in an active or inactive state. Thus, the sleeve or collar 802 can have a first dimension, such as a length when in one of the active or inactive states (e.g., Fig.58A804 ), which is greater than the second dimension of the sleeve or collar 802 when the sleeve or collar 802 is in the other of the active or inactive state. In addition, the overall length of the sleeve or collar 802 and the actuator 804 assembly or combination can be substantially constant regardless of whether the sleeve or collar 802 is in the inactive state or the active state. This difference in the size of the sleeve or collar 802 can change whether the actuator 804 is in a compressed state or at least partially uncompressed state when the sleeve or collar 802 is in the active or inactive state.

[0227] For example, according to some embodiments, Fig.58B As shown, when the shape memory sleeve or collar 802 has at least a first size or length that facilitates compressing the actuator 804 into the compressed state, the actuator 804 can be in its compressed state. Fig.58C As shown, when the sleeve or collar 802 is actuated so that the shape memory of the sleeve or collar adjusts the sleeve or collar 802 from a first size to a second, smaller size, the spring actuator 804 can extend from a compressed state to an extended or partially extended state, and thereby at least help accommodate even a temporary adjustment in the corresponding length of the powered intramedullary nail 800.

[0228] Additionally, the displacement of the actuator 804, and thus the adjustment of the relative positions of the first and second portions 810a, 810b, may be controlled using an adjustable controller 816 (eg, a pin or screw). Fig.58A In the illustrated embodiment, the adjustable controller 816 can be received in a hole 817 in the second portion 810b and in a slot 818 in the outer wall 808 of the first portion 810a, the slot 818 being sized to be received in the interior region 812 of the second portion 810b. According to such an embodiment, the controller 816 can be mechanically adjusted by the surgeon and can be configured to prevent the fractured bone segments from being pulled apart during movement of the actuator 804. Additionally, according to certain embodiments, the controller 816 can be secured to prevent relative displacement of the first and second portions 810a, 810b.

[0229] (5) Use a resorbable encapsulant to activate the balance spring.

[0230] refer to Fig.59According to some embodiments, the dynamic intramedullary nail 900 may include one or more biasing elements 902a, 902b, such as springs, each of which is encapsulated in a resorbable shell 904. The pupil of the balancing spring is to shield the fractured bone from excessive and impact forces that may interfere with the healing process. In addition, the size of the encapsulated biasing elements 902a, 902b and the intermediate element 906 can be set to be positioned in a narrow groove in the dynamic intramedullary nail 900. The narrow groove can be located at the proximal end, middle shaft or distal region of the nail. The resorbable shell 904 can prevent the activation of the biasing elements 902a, 902b. In this case, the biasing elements 902a, 902b can apply a force against the intermediate element 906 via the resorbable shell 904, such as a part of the screw positioned between the resorbable shells 904. The associated biasing elements 902a, 902b as the resorbable shell 904 can be in an active state, wherein the biasing elements 902a, 902b are extended toward and / or toward a portion of the resorbable shell 904, or in a fully expanded or relaxed state, and wherein the biasing elements 902a, 902b can only provide support for compressing the fracture gap. The use of such biasing elements 902a, 902b is advantageous in making the dynamic intramedullary nail 900 self-adjusting and controlled, and shielding the fracture from excessive and impact forces that may interfere with the healing process.

[0231] (C) Variable stiffness nail in the anterior-posterior (A / P) and medial-lateral (M / L) planes: The ability to independently control bending or torsional stiffness in the A / P and M / L planes may allow intramedullary nails to be structured to achieve optimal fracture union or reduce the incidence of periprosthetic fractures. If the patient undergoes joint replacement, a localized reduction in bending stiffness at the distal end of the nail may prevent the risk of prosthetic fracture. Fig.60 The cross section shown in Item BH in provides an example of a cross-sectional geometry for an intramedullary nail that can reduce bending stiffness in the A / P plane by approximately 20% while relatively maintaining stiffness in the orthogonal M / L plane. Again, this cross-sectional geometry can be adapted only for additive manufacturing. Fig.60 Compared to the circular cross section shown in item A, Fig.60 The various cross sections shown in Items BH can provide the following moment of inertia reductions in the A / P plane: 11% for Item B, 17-28% for Item C; 7% for Item D; 8% for Item E; 14% for Item F; 7% for Item G; and 10% for Item H.

[0232] Fig.61A and 61B Torsional and bending stiffness data based on a standard 10 mm Trigen tibial nail and a 10 mm Trigen Meta tibial nail with slots of varying sizes on the outer surface of the nail are provided. Fig.61A and61B The sensitivity of cross-sectional geometry to mechanical properties is shown. As shown, including narrow (1.5 mm), wide (3 mm), or stepped grooves (1.5-3 mm) on the outside of the nail in the A / P plane reduces the torsional stiffness from 3.5 Nm / degree to less than 1 Nm / degree compared to a standard nail. Similarly, the bending stiffness also increases from 23 Nm / degree to less than 1 Nm / degree. 2 Reduced to less than 20Nm 2 .Although Fig.61A and 61B The data shown relate to including slots located outside the nail, but the data generally highlights the principle of adjusting the nail stiffness in a specific anatomical plane.

[0233] (D) Internal geometry that provides a lower stiffness implant for larger patients: The ability to fill the intramedullary canal of larger patients with a lower stiffness implant can help accelerate the healing of fractures, especially if the bone is pathological. To avoid any complications associated with externally designed geometric features (e.g., bone ingrowth), a number of designs that exploit the design freedom of additive manufacturing are outlined below.

[0234] Fig.62 A cross-sectional view of an intramedullary nail 1000 is shown having a tapered inner wall 1002 at a mid-portion 1004 of the nail 1000 between the distal and proximal screw holes of the nail 1000. According to certain embodiments, the tapered inner wall 1002 may be dovetail shaped. Such a tapered configuration on the inner surface 1006 of the cannula of the nail 1000 may have a significant effect on torsional stiffness, which may have a beneficial effect if this design feature intersects a fracture site. In addition, such a nail 1000 may allow for rapid absorption of loads during the early stages of fracture healing without compromising the fatigue performance of the nail 1000.

[0235] Fig.63 A table is provided identifying the theoretical bending and torsional stiffness in the middle portion 1004 ("custom wall thickness") of an intramedullary nail 1000 having a tapered inner wall 1002 and an outer diameter 1008. For comparison purposes, a Trigen Meta tibial nail ("standard") is also included. Fig.63 The reduction in theoretical bending and torsional stiffness for an intramedullary nail 1000 having a tapered inner wall 1002 for the three identified outer diameter 1008 sizes, namely, outer diameters of 10 millimeters (mm), 11.5 mm, and 13 mm, assuming a constant wall thickness of 1.5 mm, is 17%, 27%, and 21.3%, respectively. The theoretical bending stiffness of a 10 millimeter (mm) outer diameter nail 1000 equipped with a customized tapered inner wall 1002 in the middle portion 1004 is calculated to be 42.5 Nm 2This custom portion of nail 1000 has a bending stiffness comparable to that of a solid profile AO universal tibial nail (Synthes) with an outer diameter of 9 mm, which was determined to be 40 Nm 2 ; Unslotted profile nail. B&K nail, which is determined to be 45Nm 2 ; and slotted profile nails with a wall thickness of 1.2 mm. K&S nails, which are determined to be 40 Nm 2 . Also like Fig.63 As shown, in torsion, the theoretical torsional stiffness of the custom nail 1000 having a tapered inner wall 1002 in the middle portion 1004, a 1.5 millimeter (mm) wall thickness, and a 10 mm outer diameter is 29.8 Nm. 2 / 2. This hardness is comparable to that of a 9.0 mm (mm) outer diameter Russell Taylor triangular nail (22.5 Nm 2 ) and 8.5 millimeter (mm) outer diameter Trigen Met tibial nail (18.4 Nm 2 )quite.

[0236] Figures 7A-7D , 64A-66 and 69 provide additional implant designs that provide reduced stiffness for larger patients. Such designs may include recesses that enter or pass through the nail wall. For example, Fig.64A A cross-sectional view of a portion of the middle portion 1102 of the intramedullary nail 1000 is shown, and Fig.64B 1000 is provided with circumferentially arranged inner grooves 1104 in its inner wall portion 1106. The grooves 1104 can be positioned in the middle portion 1102 of the nail 1100 between the distal screw hole 1108a and the proximal screw hole 1108b. Including the grooves 1104 can reduce the theoretical bending stiffness of the nail 1100. For example, a standard 10 mm outer diameter, 4.8 mm inner diameter nail with a circular cross-sectional shape can have a bending stiffness of 51.2 Nm. 2 However, including groove 1104 into such a nail can reduce the bending stiffness to 48.5 Nm 2 .

[0237] refer to Figure 7B , Fig.65 and Fig.66According to other embodiments, the dynamic intramedullary nail 1200 may include a series of circumferentially arranged open channels 1202 extending within and through the wall 1204 of the dynamic intramedullary nail 1200. The wall 1204 may include an inner wall portion 1203, which generally defines a hollow interior region 1205 of the dynamic intramedullary nail 1200. The channel 1202 may extend along a middle portion 1206 of the dynamic intramedullary nail 1200, such as in an area between the distal and proximal screw holes 1208a, 1208b. In addition, the channel 1202 may be generally parallel to a central longitudinal axis 1210 of the dynamic intramedullary nail 1200. The channel 1202 may also be arranged along one or more diameters. For example, as Fig.66 As shown, according to some embodiments, at least a portion of channel 1202 can be arranged around a first outer diameter 1212a, and other channels are arranged around a second smaller inner diameter 1212b. Channel 1202 can also have various shapes and sizes, for example, cylindrical in shape. Although the above examples have been discussed according to the middle portion 1206 of the nail, the channel 1202 structure discussed can also be used together with other parts of the implant to ensure that the fracture site can be connected with the dynamic intramedullary nail 1200 of the low modulus portion.

[0238] According to an exemplary embodiment, Fig.65 and 66 The illustrated dynamic intramedullary nail 1200 may include circumferentially arranged channels 1202 having a circular cross-sectional shape of approximately 1 mm in diameter. Including such channels 1202 in the wall 1204 of the dynamic intramedullary nail 1200 having an outer diameter of 13 mm and an inner wall portion of 4.8 mm in diameter may reduce the volume fraction of the material (e.g., titanium) in the wall 1204 from 100% to 76%, and the volume fraction of the voids in the wall 1204 provided by the channels 1202 is 24%, thereby producing a lower porosity structure. Fig.67 As shown in the table, in this case, the theoretical bending stiffness and torsional stiffness of the dynamic intramedullary nail 1200 can be increased from 156.9 Nm to 2 / 119.2Nm 2 Reduced to 110.1 / 83.4Nm 2 .

[0239] Fig.68An elastic modulus versus density plot that can be used to estimate the porosity required to match the elastic modulus of cortical bone is provided. As shown, the elastic moduli of non-porous titanium-64, 316 stainless steel, and cobalt chromium are 114 gigapascals (GPa), 193 GPa, and 235 GPa, respectively. In addition, the densities of bone, Ti-64, and 316 stainless steel can be 2.4 g / cm 3, 4.7 g / cm 3, and 8.8 g / cm 3, respectively. Using the rule of mixtures for composites, increasing the porosity of titanium-64 alloy from 0 to 40% reduces the elastic modulus from 114 GPa to 68.4 GPa, which is closer to the upper limit of cortical bone (40 GPa). Therefore, porous titanium implants can help reduce the stiffness mismatch between implants (such as intramedullary nails) and bone tissue, thereby reducing stress shielding. However, increasing porosity and pore size can lead to a reduction in the mechanical properties of the implant. Therefore, the balance between mechanical and biological properties can vary for different implant applications.

[0240] refer to Fig.69 According to another embodiment, the elastic modulus of the intramedullary nail 1300 can also be adjusted by accommodating a removable inner portion 1302 in at least a portion of the outer portion 1304 of the intramedullary nail 1300. The inner portion 1302 can be fixed to the inner surface 1306 of the intramedullary nail 1300 at one of the proximal end 1308a and the distal end 1308b of the nail, but can be locked by a locking screw (not shown) at the other of the proximal end 1308a or the distal end 1308b of the nail. As shown, the inner portion 1302 includes a wall 1310 having an outer portion 1312 and an inner portion 1314, and the inner portion 1314 of the wall 1310 generally defines a hollow interior area of ​​the inner portion 1302. The wall 1310 of the outer portion 1312 is sized to accommodate lateral displacement of the inner portion 1302 about the outer portion 1304 of the intramedullary nail 1300 and displacement of at least the connected portion of the proximal end 1308a of the intramedullary nail 1300 relative to the distal end 1308b of the nail, or vice versa.

[0241] Such a variable modulus intramedullary nail 1300 may enable the implant to exhibit a higher stiffness during fracture healing, which may be useful for at least severely comminuted fractures where greater nail stiffness and stability are desired during the initial healing period. However, the construction of the intramedullary nail 1300 may allow the elastic modulus of the intramedullary nail 1300 to be reduced in a minimally invasive manner once the fractured bone has healed, and without the need to remove the intramedullary nail 1300. For example, if an otherwise rigid intramedullary nail were removed after bone healing, the load on the bone may increase significantly, which may lead to refracture in situations with significant patient activity. Such situations may include osteoporotic fractures, where a high stiffness nail, such as a nail having approximately 300% of the bending stiffness of an osteoporotic femur, may reduce bone strength. However, after the fracture has healed, a low stiffness nail, such as a nail having a removable inner portion 1302, may be used. Fig.69 The intramedullary nail 1300 can reduce the stress shielding effect and can be relatively stiffer than a relatively stiff nail. Fig.70 The flexural and torsional stiffness of the standard Trigen Meta tibial nails with 8.5 millimeters (mm) and 10 mm outer diameters is shown to be similar to that of the Fig.13 Table 69 shows a comparison of a "custom" intramedullary nail 1300 having an outer cross-section of 10 mm. Fig.70 The "custom" nail referred to in has an inner portion 1302 having a wall 1310 including an outer portion 1312 having a diameter of 7.6 mm and an inner portion 1314 having a diameter of 4.8 mm. Fig.70 As shown, the torsional and bending stiffness of this nail can be 8.5 / 26.2Nm respectively 2 and 12.1 / 37.3Nm 2 .

[0242] Although the above-mentioned implant and method have been described in conjunction with orthopedic extramedullary nails, it should be understood that the implant and method can also be used in other technical fields and / or be associated with other types of orthopedic implants. Various changes and modifications to the described embodiments described herein will be apparent to those skilled in the art, and such changes and modifications can be made without departing from the spirit and scope of the present invention and without reducing its expected advantages. In addition, although the present invention has been shown and described in detail in the drawings and the foregoing description, it should be considered that it is illustrative rather than restrictive in nature, and it should be understood that only selected embodiments are shown and described, and it is desired to protect modifications within the scope of the present invention as described herein or as defined by the appended claims.

Claims

1. A method for manufacturing an orthopedic device, comprising: forming an additively manufactured orthopedic component from a medical grade powder and via an additive manufacturing process; heat treating the additively manufactured orthopedic component; as well as processing the heat treated additively manufactured orthopedic component to form the orthopedic device; The steps of forming the additively manufactured orthopedic component include: laser sintering the medical grade powder to form a plurality of laser sintered layers that provide a shape for the additively manufactured orthopedic component; and remelting the plurality of laser sintered layers; The remelting step includes a double scanning process along two unilateral directions; The remelting step includes alternating hatch laser grating and circumferential laser grating, wherein the alternating hatch laser grating includes scanning processes along two unilateral directions; The step of forming the additively manufactured orthopedic component includes laser sintering an outer surface and an inner surface of the additively manufactured orthopedic component, the laser sintered outer surface being separated from the laser sintered inner surface by at least a plurality of unsintered medical grade powders, and wherein the step of heat treating includes melting the plurality of unsintered medical grade powders.

2. The method according to claim 1, characterized in that The step of forming includes generating a three-dimensional model of the additively manufactured orthopedic component.

3. The method according to claim 1 or 2, characterized in that The step of forming comprises direct metal sintering using a laser power of at least 300 watts, and wherein the medical grade powder is at least a grade 5 powder.

4. The method according to claim 1 or 2, characterized in that: The medical grade powder is TiAl6V4 powder.

5. The method according to claim 1 or 2, characterized in that: The step of heat treating includes hot isostatic pressing the additively manufactured orthopedic component using a temperature of at least 1000° C. and cooling the additively manufactured orthopedic component at a cooling rate between 0.24° C. / min and 72° C. / min.

6. The method according to claim 1 or 2, characterized in that: The step of machining includes polishing the heat treated additively manufactured orthopedic component.

7. The method according to claim 1 or 2, characterized in that: The step of forming includes forming an additively manufactured orthopedic component having a non-circular cross-sectional shape.

8. The method according to claim 1 or 2, characterized in that The orthopedic device is an intramedullary nail.

9. The method according to claim 8, characterized in that The intramedullary nail has a radius of curvature between 84 cm and 156 cm.

10. The method according to claim 1 or 2, characterized in that: The step of heat treating the additively manufactured orthopedic component comprises: Purifying the environment in which the additively manufactured orthopedic component is located to a pressure of 15 mb; raising the temperature of the additively manufactured orthopedic component to between 920° C. and 1000° C.; Maintaining the pressure of the environment in which the additively manufactured orthopedic component is located at 98 MPa to 108 MPa; and The temperature of the additively manufactured orthopedic component is cooled at a rate of 10° C. / minute or less.

11. The method of claim 10, further comprising the step of stress relieving the heat treated additively manufactured orthopedic component.

12. The method of claim 10, further comprising the step of annealing the heat treated additively manufactured orthopedic component.

13. The method according to claim 1 or 2, characterized in that: The step of machining the heat treated additively manufactured orthopedic component includes removing an alpha shell layer of the additively manufactured orthopedic component using sandblasting.

14. The method according to claim 13, characterized in that The step of machining also includes the step of forming a compressive layer of residual stress on the orthopedic device.

15. The method according to claim 14, characterized in that The step of machining also includes the step of extrusion honing an inner surface of the orthopedic device.

16. The method of claim 1 or 2, further comprising the step of forming an internal sensor probe channel in the interior of the additively manufactured orthopedic component, the internal sensor probe channel not extending through the exterior of the orthopedic device.

17. The method of claim 1 or 2, further comprising the step of depositing heavy metal ions onto the orthopedic device to form an antimicrobial orthopedic device.

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