Methods for the design and manufacture of stiffness-matched implants for skeletal reconstructive surgery

Abstract

A bone fixation plate may include a first end and a second end opposite and spaced apart from the first end along a longitudinal axis. The bone fixation plate further includes a first surface extending from the first end to the second end and a second surface opposite from the first surface. The bone fixation plate further includes a network of interstices disposed between the first surface and the second surface. The plurality of interstices is provided based on a stiffness matching system and method with respect to an underlying bone or a virtual surgical model thereof. The bone fixation plate further includes a plurality of screw holes defined by and extending through the bone fixation plate from the first surface to the second surface.

Description

Attorney Docket No. 103361-579WO1T2024-046METHODS FOR THE DESIGN AND MANUFACTURE OF STIFFNESS- MATCHED IMPLANTS FOR SKELETAL RECONSTRUCTIVE SURGERYCROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This application claims priority to U.S. Provisional Patent Application No. 63 / 519,891, filed August 16, 2023, which is incorporated herein by reference in its entirety.BACKGROUND

[0002] Traditional, off-the-shelf hip, knee, spine, and craniofacial implants and fixation hardware are often successful at both holding fractures closed (Figure 1 A) and immediately restoring ambulatory, protective, or other critical functions such as speech and chewing. See, for example, FIG. 1 A, showing an example of fixation hardware preventing flexion at the fracture of a jawbone during the healing period.

[0003] However, a high percentage of surgeries are revision surgeries to fix or replace an existing implant. According to three separate studies, 36-39% of mandibular fixation devices can be expected to fail and require revision surgery. Existing devices are known to over-concentrate chewing stress within the implant itself resulting in plate fracture, screw pull-out, or both. This is often a result of stiff hardware leading to load concentrations and fracture. See, for example, FIG. IB, which shows a fragmented forearm fixation plate fracture.

[0004] Therefore, a need exists for bone fixation plates that distribute forces appropriately and minimize failure modes. The systems, methods, and devices disclosed herein relate to mandibular fixation devices as well as other skeletal reconstruction and joint replacement devices.SUMMARY

[0005] One implementation of the present disclosure is a bone fixation plate. The bone fixation plate includes a first end and a second end opposite and spaced apart from the first end along a longitudinal axis. The bone fixation further includes a first surface extending from the first end to the second end and a second surface opposite from the first surface. The bone fixation plate further includes a network of interstices disposed between the first surface and theAttorney Docket No. 103361-579WO1 T2024-046 second surface. The bone fixation plate further includes a plurality of screw holes defined by and extending through the bone fixation plate from the first surface to the second surface. The network of interstices is configured to reduce an overall stiffness of the bone fixation plate as compared to a solid reference plate, the overall stiffness substantially matching that of an underlying bone on which the bone fixation plate is implanted.

[0006] In some implementations, one or more characteristics of the network of interstices is selected based on a virtual surgical model of the underlying bone and adjacent anatomy of a patient.

[0007] In some implementations, the network of interstices includes a network of pores.

[0008] In some implementations, each interstice in the network of interstices has either (i) a uniform shape and size or (ii) a non-uniform shape and size.

[0009] In some implementations, the bone fixation plate further includes a plurality of struts extending between at least a portion of the bone fixation plate between the first surface and the second surface, the plurality of struts defining the network of interstices.

[0010] In some implementations, the plurality of struts includes a first plurality of struts extending longitudinally along the bone fixation plate and a second plurality of struts extending substantially perpendicularly to the first plurality of struts.

[0011] In some implementations, each of the interstices in the network of interstices has a substantially rectangular shape.

[0012] In some implementations, each of the interstices in the network of interstices is substantially circular.

[0013] In some implementations, the longitudinal axis of the bone fixation plate is curvilinear, and the first surface and the second surface have a curvature matching that of the longitudinal axis.

[0014] In some implementations, the curvature of the first surface and the second surface matches that of an underlying bone on which the bone fixation plate is implanted.

[0015] In some implementations, the overall stiffness of the bone fixation plate is selected based on a virtual surgical model including an underlying bone and a surrounding tissue with their corresponding mechanical properties.Attorney Docket No. 103361-579WO1T2024-046

[0016] In some implementations, the bone fixation plate includes an outer solid shell extending between the first surface and the second surface along a periphery of the bone fixation plate.

[0017] In some implementations, the plurality of screw holes are threaded and countersunk with respect to the first surface.

[0018] In some implementations, an internal porous core defined by the network of interstices includes a pore geometry to modulate stiffness, wherein the pore geometry includes orthogonally oriented struts of 300-micrometer diameter interspersed by 300-micrometer pores.

[0019] In some implementations, the bone fixation plate includes a shape memory alloy, such as a NiTi alloy (e.g., nitinol).

[0020] Another implementation of the present disclosure is a system including a bone fixation plate, at least one manufacturing device, and a control system. The bone fixation plate includes a first end and a second end opposite and spaced apart from the first end along a longitudinal axis. The bone fixation plate further includes a first surface extending from the first end to the second end and a second surface opposite from the first surface. The bone fixation plate further includes a network of interstices disposed between the first surface and the second surface. The bone fixation plate further includes a plurality of screw holes defined by and extending through the bone fixation plate from the first surface to the second surface. The at least one manufacturing device is configured to produce at least one portion of the bone fixation plate. The control system is in communication with the manufacturing device. The control system includes computer-readable instructions for producing the bone fixation plate, the computer-readable instructions being based on an anatomical model of a patient.

[0021] In some implementations, the at least one manufacturing device includes an additive manufacturing device configured to produce the bone fixation plate and the network of interstices therein according to the computer-readable instructions of the control system.

[0022] In some implementations, the additive manufacturing device produces the bone fixation plate having a desired stiffness according to the computer-readable instructions of the control system.

[0023] In some implementations, the additive manufacturing device produces the bone fixation plate further having a desired shape or a desired curvature according to the computer- readable instructions.Attorney Docket No. 103361-579WO1T2024-046

[0024] In some implementations, the at least one manufacturing device further includes a bending manufacturing device configured to adjust a shape and / or a curvature of the bone fixation plate according to the computer-readable instructions of the control system.

[0025] In some implementations, the at least one manufacturing device includes a bending manufacturing device configured to adjust a shape and / or a curvature of the bone fixation plate according to the computer-readable instructions of the control system.

[0026] In some implementations, the network of interstices is fabricated, via the manufacturing device, to have a pattern matching the desired stiffness of the bone fixation plate.

[0027] In some implementations, the manufacturing device is a metal-powder-based 3D printing device.

[0028] In some implementations, the network of interstices includes a network of pores, each of the pores in the network of pores having either a uniform shape and size or a non-uniform shape and size.

[0029] In some implementations, the bone fixation plate further includes a plurality of struts extending between at least a portion of the bone fixation plate between the first surface and the second surface, the plurality of struts defining the network of interstices.

[0030] Another implementation of the present disclosure is a method of fabricating a bone fixation plate. The method includes: providing an anatomical model of a patient; calculating a stiffness of a portion of the anatomical model and the curvature of a portion of the anatomical model; calculating a set of computer-readable instructions for producing the bone fixation plate having a shape matching the stiffness and the curvature of the portion of the anatomical model; and fabricating, via a manufacturing device configured to receive the set of computer-readable instructions, the bone fixation plate with a network of interstices providing a stiffness of the bone fixation plate that substantially matches the stiffness of the portion of the anatomical model.

[0031] In some implementations, the manufacturing device is an additive manufacturing device including a metal-powder material.

[0032] In some implementations, a plurality of struts define the network of interstices of the bone fixation plate, the plurality of struts being arranged orthogonally to each other.Attorney Docket No. 103361-579WO1 T2024-046

[0033] In some implementations, the method further includes implanting the bone fixation plate on an anatomical feature of the patient on which the anatomical model was based by installing a screw through a screw hole extending through the bone fixation plate.

[0034] In some implementations, the bone fixation plate is configured to provide sufficient support to a portion of a patient's anatomy and to provide sufficient flexibility to reduce stress fractures of the bone fixation plate.

[0035] The systems, methods, and devices are explained in even greater detail in the following drawings. The drawings are merely exemplary and certain features may be used singularly or in combination with other features. The drawings are not necessarily drawn to scale. Any dimensions shown in the drawings are exemplary only and do not limit the scope of the disclosure.

[0036] Additional advantages will be set forth in part in the description which follows or may be learned by practice. The advantages will be realized and attained by means of the elements and combinations particularly pointed out in the appended claims. It is to be understood that both the foregoing general description and the following detailed description are exemplary and explanatory only and are not restrictive, as claimed.BRIEF DESCRIPTION OF THE DRAWINGS

[0037] FIG. 1 A shows a craniofacial fracture with implanted hardware showing the associated forces on the implant, according to one implementation.

[0038] FIG. IB shows an image of a fragmented forearm fixation plate fracture, according to one implementation.

[0039] FIG. 2 shows a diagram of a workflow for manufacturing a device or implant, according to one implementation.

[0040] FIG. 3 shows a workflow for designing a patient-specific mandibular fixation plate, according to one implementation.

[0041] FIG. 4A shows a series of prototype, 3D printed, porous bone fixation plates, according to one implementation.

[0042] FIG. 4B shows prototypical porous bone fixation plates on bone, according to one implementation.Attorney Docket No. 103361-579WO1 T2024-046

[0043] FIG. 5A shows a diagram of a method for facilitating the uploading of a patient's 3D CT scan and estimating the stiffness required for a bone fixation plate, according to one implementation.

[0044] FIG. 5B shows a 3D structure of the example mandible of FIG. 5 A with cross-sectional views along various planes, according to one implementation.

[0045] FIG. 5C shows the 3D model of FIG. 5B with a layout for the location of fixation hardware, according to one implementation.

[0046] FIG. 5D shows a visualization graph showing the stress (Von-Mises Stress) experienced by the fixation hardware once in place on the 3D model of the mandible, according to one implementation.

[0047] FIG. 6 shows a diagram of the process of creating a porous fixation device, according to one implementation.

[0048] FIG. 7A shows various views of a bone fixation plate, according to one implementation.

[0049] FIG. 7B shows various views of a bone fixation plate used in the described study, shown with exemplary dimensions, according to one implementation.

[0050] FIG. 8 shows a screw planning finite element model (FEM), according to one implementation.

[0051] FIG. 9 shows a schematic illustration of optimization of the stiffness-matched orthopedic device and virtual surgical planning, according to one implementation.

[0052] FIG. 10 shows a diagram of the basic approaches to a stiffness-matching objective, according to one implementation.

[0053] FIG. 11 shows a graph of modulus of elasticity of biomedical alloys, according to various implementations.

[0054] FIG. 12 shows a schematic illustration of the 2-D finite element model and spatial variation of the maximal shear stress in the periprosthetic bone tissue for different implant roughness and bone-implant contact ratios, according to one implementation.

[0055] FIG. 13 shows, in panel A, CT-Renderings of bone ingrowth into a locally stiffness- matched porous scaffold, in panel B, fluorescent microscopy images with islets of bone deepAttorney Docket No. 103361-579WO1 T2024-046 within the scaffold interior circled, and, in panel C, histology images with magnified sections of the scaffold periphery and interior respectively, according to experimental implementations.

[0056] FIG. 14 shows a diagram of test set-ups, example prototypes, and displacement / compression diagrams for hybrid meta-biomaterials, according to various implementations.

[0057] FIG. 15 shows a diagram illustrating calculation and modeling of gradient porosity to yield desired mechanical properties for bone implants, according to various implementations.

[0058] Various objects, aspects, features, and advantages of the disclosure will become more apparent and better understood by referring to the detailed description taken in conjunction with the accompanying drawings, in which like reference characters identify corresponding elements throughout. In the drawings, like reference numbers generally indicate identical, functionally similar, and / or structurally similar elements.DETAILED DESCRIPTION

[0059] Referring generally to the figures, stiffness-matched bone fixation plates, and a method of manufacturing the same, are shown, according to various implementations.

[0060] There are primarily two types of skeletal reconstructive devices, skeletal fixation, and skeletal replacement devices. These devices can be uni-material or multi-material and fabricated by many or a single method (e.g., 3D printing, milling, grinding, molding, deformation, etc.). The systems, methods, and devices disclosed herein provide a unique approach to the design and manufacture of these devices.

[0061] The systems, methods, and devices disclosed herein can be designed to be either generic or personalized to a particular patient. The approaches described herein include the design of the device's shape and materials for use in computer-based and possibly physical (real- world) simulations of optimal device performance. Performance can be optimized by iteratively or algorithmically modifying: (a) device geometry, (b) device location, and (c) device material. A "stiffness matching" approach tunes the device's mechanical properties to accomplish the job at hand (e.g., replacing / correcting bones, bone portions, bone grafts, ligaments, tendons, and any combination of these) without stress shielding healing segments of bone, creating stress concentrations in a device that cannot sustain them for as long as is needed, or interrupting normal stress-strain trajectories to the extent that the healed (repaired) bones do not have a sustainable loading pattern (i.e., one that involves sustainable remodeling in response toAttorney Docket No. 103361-579WO1T2024-046 sustainable load). The design of the device shape, materials, and optimal performance also includes a plan for fabrication where the fabrication steps only reinforce and do not disrupt, the restoration of the intended function.

[0062] Devices prepared using the methods described herein may be optimized over current standard-of-care devices to obtain: (1) reduced stiffness allowing better engagement of the whole plate, avoiding stress concentrations; (2) better engagement of the whole plate and accurate screw planning reducing screw loosening (i.e., instead of a screw that is adjacent to a heavily loaded screw being loosened, it is tightened); (3) tightening is enhanced through accurate bicortical and changing screw orientation (i.e., not all in the same plane) when the plate wraps around the bone; and (4) carefully allowing the plate to wrap around the bone helps keep the host and graft bone fragments in compression during loading, thus speeding the healing process.Hypothetical Framework for Robotic Manufacturing of Mandibular Graft Fixation Plates

[0063] A hypothetical workflow for manufacturing a disclosed device or implant includes four stages, as shown in FIG. 2. Specifically, FIG. 2 shows a diagram of a hypothetical robotic Point- of-Care Manufacturing (POCM) workflow including Stage 1 : Segmentation (identification) of patient 3D CT surfaces of interest; Stage 2: Implant Design and Mechanical Modeling: Validation of plate location, fit, and screw paths; Stage 3: Manufacturing Process Engineering coded robot tool paths; Stage 4: Bending, Twisting, and Peening leads to a flush-fitting fixation plate.

[0064] The process starts with the CT scan of the region of interest and the segmentation of the anatomical surfaces to be reconstructed. Then, the bone model is processed in a Virtual Surgical Planning (VSP) environment as in surgery (i.e., cut, reconstructed, engrafted, etc.). In this virtual environment, it would be possible to design a personalized implant to fit the original or reconstructed anatomy. Afterward, the mechanical performance of the reconstructed, and fixated, bone graft would be computationally assessed (i.e., applying a static load) via Finite Element Analysis (FEA) and further optimized to enhance the surgical outcome. This workflow includes the personalization of a medical device based on the mechanical requirements. Finally, the manufacturing process would be based on metal forming strategies, which would be previously validated via process engineering, to ensure personalized devices with desired mechanical properties as an outcome.Attorney Docket No. 103361-579WO1 T2024-046

[0065] It is important to mention that in these stages the study considered the use of a quality management system (QMS) not only for the fixation plates that would be manufactured but also for the hardware and software considered in the workflow and the manufacturing process. In this sense, design and fabrication processes would be documented to maintain their effectiveness according to the requirements of international standards, such as ISO 13485:2016 (i.e., Medical Device Good Manufacturing Process) for medical devices.Stage 2 (Medical Device Shape and Mechanical Personalization)

[0066] Patient-specific mandibular graft fixation may be designed in Geomagic Freeform (3D Systems, Rock Hill, SC, US) software. The software allows the detection of the mandibular surface. A generic straight fixation plate design may be used to create a bent plate that is fully in contact (flush) with the mandibular and graft surfaces. The Geomagic digital tools also facilitate a virtual surgery simulation that considers the location of the mandibular resection, geometry, and length of the bone graft during implant design. FIG. 3 shows a workflow for designing a patient-specific mandibular fixation plate. A line, which would be the undersurface centerline of the implant, is drawn over the bone surface (e.g., on the mandibular and bone-graft (i.e., osteotomy) surfaces which represents the midline of the desired fixation plate’s location). Next, the implant design is warped based on that line. A series of operations would be used to create the personalized plate with the desired external dimensions (i.e. length, width, and thickness) based on the mechanical needs of the healing process as well as consideration towards not interrupting future normal loading of the healed bone. Then, the mechanical modeling includes planning screw and screw hole location and length. Finally, cutting guides would be designed. If useful, the host mandible, bone graft, screws, and implant could be exported for testing and optimization, by computational and / or in-vitro mechanical analysis, of the overall reconstruction’s mechanical performance during mock chewing. The main purpose of the fixation plate is to hold in close contact the graft bone with the host mandible and to offer stability to the graft union.

[0067] The mechanical behavior and strength of the fixation plate, fixated bone graft, and host mandible would be simulated during mastication via static FEA for two scenarios of interest: (1) during the healing period to evaluate the implant’s stiffness and stability, and (2) after bone healing and muscle force restoration is complete to avoid stress shielding of that newly healed bone. To this end, a volume mesh of the 3D CT or CAD-derived components (fixation plate, host bone, screws, and bone graft) would be created, the boundary conditions (displacementAttorney Docket No. 103361-579WO1 T2024-046 restraints and forces) would be set, the material properties would be selected, the interaction between components would be established, and the model would be solved. Furthermore, preliminary mesh quality and mesh convergence studies may be performed to increase the accuracy of the FEA results. Boundary conditions simulate chewing for maximum occlusion at the right first molar (Ml) by restraining the movement in all directions of the buccal cusps of the teeth when they are inside the two rows of upper cusps (i.e., centric occlusion). The mandibular condyles would be constrained to prevent movement as well. Each masticatory muscle's force magnitude, direction, and area of attachment would be defined. However, 60% of the maximum value would be used during the before-healing computational analysis as chewing power decreases after mandibular reconstructive surgery and is slowly regained.

[0068] The contact between the host bone and the grafted bone would be simulated as well for two scenarios: before healing (no union) and after healing (union). Others have analyzed the prehealing state of engrafted bone by assigning a friction coefficient of zero between components to allow free motion. After computational analysis, the host bone-graft bone interface micromotion and reaction force, as well as the resulting Von-Mises stress distribution in the bone and implant, would give feedback to the plate design stage for implant optimization. With these results, it could be assured that the bone graft is in compression and the maximum micromotion value (300-400 um) is reached, both critical for successful healing. Additionally, the stress distribution results would show the location of stress concentrations and thereby potential areas of failure to optimize (remove) in the design of the skeletal fixation plate. Thus, the process of iterative design of the plate, screw depth, location planning, and validation by mechanical testing would ensure before implantation that the performance-optimized plate was obtained by the optimized POCM process.

[0069] This pre-operative mechanical model of chewing could be used to interactively change the size, shape, or location of the fixation device, which has been demonstrated to have an impact on the reduction of stress shielding in implants. These variations are all done to accomplish three things simultaneously: optimal healing outcome; post-healing lack of stress shielding; and fabrication process engineering designed to achieve both the personalized shape and mechanical function of the fixation plate. In the ideal situation, the patient’s surgeon would have input into these decisions. That rarely occurs in current commercially available service workflows. When personalized plates are ordered from commercial vendors, the physician’s input may be limited to approval of the device’s final shape for delivery.Attorney Docket No. 103361-579WO1T2024-046Stage 3 (Process Engineering)

[0070] Once the optimized design of the fixation plate is obtained, manufacturing planning can proceed (e.g., process engineering, see FIG. 2). At this point, the curvature ranges and twisting angles of the plate would be determined from the optimized design and mechanical simulation performed in stage 2. To fabricate the fixation plates, the deformation strategy to obtain the primary shape would be performed by roll bending. Inspired by the performance of automated and flexible metal forming techniques, specific deformations or twisting may be performed on the plate with a robotic system to achieve the shape determined in the previous stage.

[0071] This back-and-forth stage would also serve to validate the optimization of the fixation plate’s performance and optimization of the fabrication process to produce a fixation plate with that performance. This would be accomplished by using an ICME validation model to help predict the microstructural evolution of the plate's material based on the design variables and forming process parameters. This data would also help determine, through computational simulations of the metal forming process, the forming loads, spring back, or specialized fixturing for the available plate bending equipment. The simulation and validation of robot trajectories and forces for the fabrication of fixation plates would be translated into Robot Operating System (ROS) process controls. This would allow a digital twin to validate the manufacturing of the plates obtained from the design stage.

[0072] One example of a Stage 3 device, method, and system is shown in the below-described Example device, shown in FIG. 7 A, with additional examples in FIGS. 4A-4B and FIG. 7B. The systems and devices shown in the figures (e.g., in FIGS. 4A-4B and 7A-7B) can accept and read instructions based on the above-described steps (e.g., Stage 1 and Stage 2 data) and perform operations to shape the fixation plate to match a patient’s anatomy. The systems and devices may further implement additional features and structures elsewhere described herein (e.g., after- treatment operations, metal hardening, and / or other microstructure-related operations).Stage 4 (POCM: Fabrication Modalities and Sequencing for Deformation of Graft Fixation Plate)

[0073] This stage consists of 4 sequential strategies based on a Hybrid Autonomous Manufacturing-Metamorphic Manufacturing (HAM-MM) approach, that would produce the final fixation plate to design specifications and consider the forming loads determined in the previous stage. The first step would consider bending straight plates by employing a slip rollerAttorney Docket No. 103361-579WO1 T2024-046 of varying diameters according to the curvature ranges determined in the design phase. Then the plate would be delivered to a station and fixed in a vise press, so the robotic system can apply the determined loads and angles to twist it. The final rough tuning of the surface contacting the bone (to adapt the surface of the plate to that of the mandible) would be made by peening. This operation should minimize the space between the two surfaces for proper fixation. The eyelet for fixation screws would be threaded (i.e., either standard threading or locking head threading) using 5-axis CNC machining. The location of each hole is determined by the previous Virtual Surgical Planning (VSP). In addition, air gas will be used for chip removal during the process to prevent the cutting tool or plate from cracking due to material entrapment. Finally, the ends of the plate would be cut and polished to achieve the final geometry. It is important to mention that sensory systems and control algorithms (MM approach) would be used to track the fabrication process to ensure that the plates better adhere to quality standards and performance requirements identified in the VSP stage.Framework for Stiffness-Matching for Manufacturing of Mandibular Graft Fixation Plates

[0074] To address issues with existing devices, an implant design and manufacturing methodology for nitinol (NiTi aka nickel -titanium) skeletal fixation hardware was created that is stiffness-matched to an individual’s needs. Porous Nitinol devices have been developed with the same stiffness as, or even less stiff than, bone. In this way, the problems of stress shielding and stress concentration that cause current skeletal fixation devices to fail may be solved. Such devices may be fabricated via 3D printing. Such devices may include both porosity and superelasticity, a unique property of Nitinol (NiTi), to match stiffness to local needs to ensure the regeneration of reconstructed bone and surrounding muscles.

[0075] The developed devices provide short-term skeletal fixation and healing, avoid longterm functional problems, and provide a manufacturing advantage over the current competition. For example, current standard-of-care skeletal fixation hardware fails because it is several times stiffer than any human being’s bone. The current failure rate, for example, nearly 40% with mandibular skeletal fixation (FIG. 1 A), followed by required revision surgery, is unnecessary.

[0076] As the first example of this technology, it has been demonstrated that it is possible to produce mandibular (lower jaw) Nitinol fixation hardware with the appropriate stiffness for reconstructive surgery due to trauma, post-cancer resection, sports or aging injury, or congenital deformity. Prototypes have been successfully 3D printed, as shown in FIG. 4A standing alongAttorney Docket No. 103361-579WO1 T2024-046 and in FIG. 4B placed on bone, and mechanically tested to demonstrate that each relevant stiffness level can be produced.

[0077] An alpha version of end-user (i.e., physician’s office and Radiology Department) software has been developed that facilitates the uploading of a patient’s 3D CT scan as shown in FIG. 5A. Such software may be implemented into the framework shown and described in FIG. 2 (e.g., in stage 2 for VSP design and stage 3 for process engineering). The software may ask surgeons to identify key structures (e.g. bones, muscles, etc.) to construct a functional model of the structures to be repaired. In the mid and lower face, the treatment plan allows for the design of fixation hardware that would re-establish the normative load distribution of these structures, which is greatest during chewing.

[0078] In this portion of the disclosed method, as shown in FIG. 5A, the first step for the design of patient-specific fixation hardware is obtaining 3D CT-scan data to create a CAD file of the mandible. The 3D CT-scan data are again used to estimate the required stiffness for the fixation hardware. The result is a biomechanical model of the mandible that includes properties of the tissues (i.e., cortical bone, cancellous bone, periodontal ligaments, teeth, and masticatory muscles) as well as the muscle forces. See also FIGS. 5B-5D showing additional steps and images from the design and planning process. Specifically, FIG. 5B shows a 3D structure of the example mandible with cross-sectional views along various planes. FIG. 5C shows the 3D model with a layout for the location of fixation hardware. FIG. 5D shows a visualization graph showing the stress (Von-Mises Stress) experienced by the fixation hardware once in place on the 3D model of the mandible.

[0079] The CAD files of the host mandible (healthy regions of the patient’s mandible), double barrel fibula graft (or sometimes single barrel due to the limitations or special conditions) to fill a defect, and any other required filling elements (e.g., soft tissue or an inert space holder) are assembled to simulate the reconstruction surgery. The surgical planning software allows the treatment team to assess the variation in the design of the fixation hardware to achieve a long- lasting treatment. This is achieved by restoring the normal distribution of stress after the surgery. The team can also decide on the best location for plates and screws to obtain the desired result.Adding porosity to achieve long-lasting fixation hardware

[0080] To achieve long-lasting mandibular reconstruction the mechanical properties of the fixation hardware may be tuned to restore the normal stress distribution after surgery. Example surgical simulations and treatment planning software have been developed to simulate fixationAttorney Docket No. 103361-579WO1 T2024-046 devices with varying porosity and pore geometry on the long-term health of the reconstructed mandible.

[0081] The results of simulation, fabrication, experimental evaluation, and biocompatibility assessment indicate the technology readiness of additively manufactured Nitinol fixation hardware to achieve long-term success in reconstructive surgery of the mandible. See the diagram of FIG. 6 showing the process of creating a porous fixation device, according to one implementation. The process includes using a CAD model (e.g., the model derived from the patient-specific model used for visual surgical planning) and laying out one or more porosity graphs to generally match the stiffness data and requirements for that patient and / or implant. Then, thicknesses may be assigned and adjusted (e.g., for the struts or pores) to further adjust the stiffness. Finally, a shell may be added to the outer surface(s) of the implant (e.g., on sides facing away from the underlying bone).

[0082] The fixation devices disclosed herein for musculoskeletal reconstructive surgery are more porous and less stiff than fixation tools currently on the market. Existing tools are harder than the underlying bone, creating stress and fracturing in healthy bone over time. The disclosed devices provide the needed flexibility to secure the joint and allow movement, without putting unsustainable pressure on bone.Example Device and System

[0083] FIG. 7 A shows a bone fixation plate 100, according to one implementation of the present disclosure. The bone fixation plate 100 includes a network of interstices, or spaces, defined in the body of the bone fixation plate. As shown, the network of interstices includes a plurality of pores 120 having a uniform, square shape and extending substantially perpendicular to the first and second surfaces of the bone fixation plate. However, in other implementations, the network of interstices may include a different arrangement of interstitial, spaces, pores, or differently shaped structures reducing the overall density and stiffness of at least a portion of the bone fixation plate.

[0084] The bone fixation plate 100 includes a plurality of struts 122 defining the plurality of pores 120. The bone fixation plate 100 further includes a first end 102 and a second end 104 opposite and spaced apart from the first end 102 along a longitudinal axis 106. The bone fixation plate 100 further includes a first surface 108 extending from the first end 102 to the second end 104. The first surface 108 is substantially flat across the entirety of the bone fixation plate 100. The first surface 108 is configured to face away from an underlying bone or other anatomicalAttorney Docket No. 103361-579WO1 T2024-046 feature. The overall shape and curvature of the bone fixation plate are configured to match the curvature of the underlying bone or other anatomical features (e.g., as derived from a virtual surgical model).

[0085] The bone fixation plate 100 further includes a second surface 110 opposite from the first surface 108. A side wall 112 (e.g., a portion of an “outer shell”) extends between the first surface 108 and the second surface 110. The side wall 112 extends around the periphery of the bone fixation plate 100 and may have a thickness greater than that of other portions of the bone fixation plate 100. For example, the side wall may be thicker than the plurality of struts 122 extending through the interior of the body of the bone fixation plate 100. The second surface 110 is configured to face towards the underlying bone or other anatomical feature. The second surface 110 is exposed to leave open a plurality of pores 120 defined in the bone fixation plate 100, as further described below.

[0086] The plurality of struts 122 extends between the first surface 108 and the second surface 110. The plurality of struts 122 defines the plurality of pores 120 that extend a distance from the second surface 110 towards the first surface 108. The plurality of pores 120 extends in a direction substantially perpendicular to each of the first surface 108 and the second surface 110.

[0087] The plurality of struts 122 includes a first set of struts 122a extending in a substantially longitudinal direction and a second set of struts 122b extending in a direction along the “width” of the bone fixation plate 100 from one portion of the side wall 112 to another portion of the side wall 112 opposite of the longitudinal axis 106. In the bone fixation plate 100, the first set of struts 122a and the second set of struts 122b are substantially perpendicular to and equally spaced from each other. Thus, the plurality of pores 120 are square-shaped. Generally, the plurality of pores 120 is configured to reduce the overall stiffness of the bone fixation plate 100 as compared to a completely solid reference bone fixation plate of similar geometry and proportions (e.g., as derived from a virtual surgical model).

[0088] The network of interstices formed by the plurality of pores 120 forms an internal porous core of the bone fixation plate 100. The network of interstices and the internal porous core formed thereby includes a pore geometry used to modulate stiffness in different portions of the bone fixation plate. For example, the pore geometry may include that each of the plurality of struts 122 is 300 micrometers thick such that the plurality of pores are 300 micrometers across in each direction. In other implementations, the struts and pores have a different size and / or percentage of porosity.Attorney Docket No. 103361-579WO1 T2024-046

[0089] In general, the size of the interstices can vary significantly between devices of different sizes and different applications (e.g., different anatomical locations undergoing different levels of loading). Biomechanical properties are influenced by the mechanical properties of the material and the geometry of the device as well as its location (i.e., where is it attached to the skeleton). For mandibular plates as shown, a relatively high load is experienced, so the struts and pores had a similar sizing (e.g., 50% porosity or air space). Similar locations in the mandible (lower jaw) or maxiallary (upper jaw) may experience similar stresses and require similar bone fixation plates with similar porosity. However, this porosity level could be dramatically different in an area under higher or lower load regimes.

[0090] In some examples, the first portion of the bone fixation plate (e.g., closer to the first end or closer to the first surface) includes interstices of a larger diameter or size and a second portion of the bone fixation plate (e.g., closer to the second end or closer to the second surface) includes interstices of a smaller diameter or size. In such an example, the first portion of the bone fixation plate would have a relatively lower stiffness than the second portion of the bone fixation plate.

[0091] However, in other implementations, the plurality of struts may extend at a different distance from each other (e.g., forming rectangular-shaped interstices). In other implementations, the plurality of struts may extend at a different angle with respect to each other (e.g., forming trapezoidal or triangular-shaped interstices). In other implementations, the plurality of struts may include three or more sets of struts. The network of interstices, or pores, may have a variety of adjustable characteristics each of which may be adjustable based on a virtual surgical model. The characteristics of the network can include, among other things, the total number of pores, the size of pores, the homogeneity of pore size, and the distribution of the network of pores throughout the bone fixation plate. In other implementations, the network of interstices may have a rounded shape, a circular shape, or an irregular shape. In other implementations, the network of interstices may include regularly and uniformly shaped and spaced interstices. In other implementations, the network of interstates may include irregularly and / or non-uniformly spaced and / or shaped interstices.

[0092] The bone fixation plate 100 further includes a plurality of screw holes 130 extending through the bone fixation plate 100 from the first surface 108 to the second surface 110. The plurality of screw holes 130 extends through the plurality of struts 122 such that a reinforced area 132 surrounds each of the plurality of screw holes 130. The reinforced area 132 includes aAttorney Docket No. 103361-579WO1 T2024-046 thicker wall than the adjacent elements and provides structure and support to the plurality of screw holes 130. Therefore, the plurality of struts 122 extends between the side wall 112 on the periphery and the reinforced area 132 on the interior of the bone fixation plate 100.

[0093] In some implementations, the plurality of screw holes may be threaded screw holes. In some implementations, the first surface includes a countersink or other sunken feature sized to accommodate a head of a screw of the plurality of screws. In some implementations, a system is disclosed, the system including the bone fixation plate 100 and a plurality of screws extending through the plurality of screw holes 130.

[0094] In some implementations, a system is disclosed, the system including the bone fixation plate 100 and an additive manufacturing device (e.g., a metal-powder-based 3D printing device) configured to produce the bone fixation plate 100. In other implementations, the network of interstices may include a porous metal or a foam metal material. In other implementations, the network of interstices are formed by laser sintering, laser cutting, laser etching, punching, or other manufacturing operations capable of forming a network of interstices in the bone fixation plate.

[0095] In some implementations, the system includes a control system in communication with the additive manufacturing device. The control system may include computer-readable instructions for producing the bone fixation plate to a desired shape and a desired stiffness, the computer-readable instructions being based on a visual surgical model of a patient (e.g., as described in reference to FIG. 2 and FIG. 5 A). For example, the computer-readable instructions may include instructions for producing the network of interstices with a particular shape, size, pattern, arrangement, and / or density, such features affecting the overall density and stiffness of the bone fixation plate.Experimental Study, Results, and Discussion

[0096] Manually bent, standard-of-care, Surgical Grade 5 Titanium Alloy (Ti-6A1-4V), off- the-shelf mandibular graft fixation devices are associated with a significant post-operative failure rate. In all cases, these failures require the patient to endure stressful and expensive reoperation. This disclosure proposes a stiffness-matched approach that facilitates physician assessment of anticipated, post-operative performance by optimizing graft fixation device location, shape, and material properties. In some implementations, the optimal performance of a device combines translation of all loading into compression of the bone graft with the adjacentAttorney Docket No. 103361-579WO1T2024-046 bone segments and elimination or minimization of post-healing interruption of normal stressstrain (loading) trajectories.

[0097] Described herein is a sheep mandibular graft model where four animals received virtually optimized, experimental NiTi (nickel -titanium) fixation plates fabricated using Laser Powder Bed Fusion 3D metal printing. A 1.75 cm mandibular graft healed completely in all four animals receiving experimental devices. The last animal received a Ti-6A1-4V standard-of-care, manually bent, fixation plate that very shortly after implantation fractured and dislocated. The disclosed design technique, especially screw orientation and depth planning improved throughout the study leading to the most rapid healing. Highlights of the study include: (i) Current standard-of-care fixation devices are 4-6 times stiffer than cortical bone; (ii) Highly stiff fixation results in bone stress shielding and resorption (bone loss); (iii) Overly stiff fixation results in device stress concentration and device failure; and (iv) Designed NiTi fixation plate porosity and super elasticity avoid these phenomena.Introduction

[0098] Reconstructive craniomaxillofacial (CMF) surgeries are commonly simulated on a computer in a Virtual Surgical Plan (VSP). In most VSP environments one can obtain the target shape of a fracture or a segmental defect site that is to be reconstructed as well as identify sites for fixation hardware placement. The size and shape of a bone graft to fill segmental defects can also be visualized. However, current surgical simulation sessions rarely provide the surgeon with the biomechanical properties of the fixation to be used or its anticipated post-operative mechanical performance. Instead, surgeon decisions on fixation hardware geometry, material, or location are most commonly made from experience. Currently, it is not infrequent that a model of the virtually reconstructed bone is 3D printed, and fixation plates are manually bent to that model to provide a fit. As an alternative to a manually bent fixation plate, it is now common to 3D print larger, VSP-designed, CMF fixation plates. This level of personalization ensures that the plate will be uniformly flush with the bone, thereby increasing the overall reconstruction’s stability,

[0099] As mentioned, beyond flush plating, predicting functional outcomes is left to the surgeon’s experience and medical judgment. Because there is virtually no biomechanical information available in current fixation design specifications provided to the physician, FDA clearance (e.g., testing conformance with standard-of-care device guidelines) panels use generic mechanical testing standards, e.g., ASTM F382-17, as a basis to reduce the risk of fixationAttorney Docket No. 103361-579WO1 T2024-046 device mechanical failure. Often, when it occurs, fixation device failure is attributed to overly stiff materials that may have led to stress shielding-induced bone resorption (loss), implant loosening, and eventually device failure. The lack of biomechanical data makes it difficult for attending physicians to avoid stress shielding or, another possible cause of failure, stress concentrations. Stress concentration, especially in work-hardened “crimp” (z.e., thinned areas for manual bending) zones, may also lead to device failure. Indeed, 36-39% of large-mandibular graft fixation devices can be expected to fail and require revision surgery, with 8-10% of all CMF fixation plates having been observed to break, loosen, or in other ways fail during normal activities (i.e., no trauma). In addition to the painful and stressful emergency caused by the unexpected failure of these devices, typical re-operation costs average $50,000.

[0100] If biomechanical data were available during VSP, it could be used by the surgeon to choose the fixation plate shape (i.e., external shape, internal porosity, screw location, type, and length), fixation plate location, and fixation plate material. Indeed, extending current VSP approaches to include this information, and to fabricate a personalized mandibular graft fixation plate based on that plan, is the direction taken in the study reported here. Good fit and function of fixation hardware are expected to improve and speed healing, which may be aided by the postoperative wound healing environment. That is because bone fractures, by virtue of hematoma formation from injured vessels, the early inflammatory response, and the cytokine cascade that follows — the surgical “injury” (e.g., soft tissue incisions, osteotomies, etc.) — start the healing process by creating an interplay between bone, endothelial, and inflammatory cells, that ultimately can aid in bone regeneration and remodeling.Materials and Methods

[0101] VSP: Skeletal Geometry and Biomechanical Input into Fixation Shape and LocationDecisions - A VSP biomechanical model that includes the host bone surrounding a defect, a bone graft filling the defect, a graft fixation device, and the anticipated chewing forces facilitate the selection of fixation hardware shape, location, and material properties that can be expected to help avoid harmful stress shielding of the bone and stress concentrations in the device. To create this model, an outer surface image of the skull and mandible, derived from a 3D CT image, is converted to a solid object composed of a tetrahedral mesh with assigned mechanical properties (i.e., a finite element model). The geometry and mechanical properties of the masticatory anatomy that is to be surgically reconstructed can be directly derived from 3D CT data. The goal is to design personalized fixation hardware that facilitates bone graft and adjacent host boneAttorney Docket No. 103361-579WO1 T2024-046 healing while not interrupting, or minimally interrupting, normal masticatory stress-strain trajectories once the bone has healed. Informing surgeons’ choices about a personalized fixation device’s shape, material properties, and location helps to avoid both stress shielding-induced bone loss and stress concentration-induced fixation device failure.

[0102] VSP: Biomechanical Input into Material Choice and Shape Modulation - Methods for the creation of a personalized model of mastication based on a 3D CT image have been previously presented. This task begins with obtaining masticatory muscle vectors (i.e., the direction of pull) and force magnitude (i.e., chewing strength). Force magnitude is directly inferred from the maximum cross-sectional area of each muscle. Next, a model for maximum occlusal force is simulated based on biomechanical data. The spectrum of fixation options, from relatively thin, simple dog-bone shaped miniplates, to L, U, square, or complex shapes, including thick reconstruction bars that are manually bent or 3D printed to fixate the mandibular graft are surveyed as are the available screw and screw hole options, including locking screws i.e., threads on screw head). Stress concentrations occurring in thin areas of a fixation device or a gap between the plate and the bone surface are avoided.

[0103] By modeling normal (i.e., healed bone) stress-strain trajectories, the normal loading patterns will be as uninterrupted as possible by the presence of CMF fixation which has a different role to play. The role of fixation is to hold the bone graft in place while it heals with the adjacent host bone. Unlike bone replacement hardware (e.g., hip, knee, spine, shoulder), fixation may be stronger than the adjacent bone. However, there is a benefit to having sufficient elasticity to engage as much of the whole plate, and as many of the screws as needed to keep the graft and adjacent bone in compression, with minimal micromotion at the osteotomy sites, under all loads. It is known that micromotion at the osteotomy (healing) sites may be kept under 300-400 mm.

[0104] The most commonly used fixation material is Surgical Grade 5 titanium alloy, Ti-6A1- 4V. This material has a much higher stiffness (112 GPa) than the surrounding mandibular cortical bone (10-31 GPa). Dense NiTi has a stiffness of between 28-70 GPa (austenite: 70 GPa; martensite: 28-40 GPa), which allows the design of porous spaces to modulate stiffness with much less risk of stress concentration failure versus Ti-6A1-4V. In addition to in vitro cytotoxicity studies, it was recently demonstrated in rat cranial and vertebral skeletal repair models that medical grade NiTi (nickel -titanium), is biocompatible and safe for use in skeletal reconstruction hardware.Attorney Docket No. 103361-579WO1T2024-046Sheep Model: 5 Polled Dorset Sheep

[0105] In this study, an example model of human mastication was applied to a sheep model for mandibular graft fixation. The example model takes advantage of the fact that the diastema region of the sheep mandible can be removed and immediately placed back without removing teeth, as would be required with most other relevant large mammal models. While the ossification centers of the sheep mandible are closed before birth, growth continues through the period when dental maturity is obtained. In addition, it allows the dimensions of the fixation hardware to be similar to that used in humans. Bone mineral composition of the sheep mandible and its rate of bone healing do not differ significantly from humans. Five (4 experimental, 1 control) Polled Dorset strain, a strain originally developed at North Carolina State University, were obtained from The Ohio State University (OSU) herd in Wooster, OH were the animal subjects of this study. Polled sheep have no horns. Because there is more variation between ewes and wethers than between Dorset sheep and other strains, only females were used. Moreover, dental measurements were used to verify the morphological similarity of the sheep chosen from the OSU Dorset herd. Others have used a sheep mandibular graft model to study the regeneration of load-bearing, skeletal, and segmental defects (i.e., the osteotomized bone is replaced with a tissue-engineered artificial graft).

[0106] Sheep Model: Design and Fabrication of Personalized (shape, location, material properties), 3D printed NiTi Mandibular Graft Fixation Plate Performance - Using the previously described biomechanical model of chewing, (VSP) maximum occlusion was virtually simulated at the first molar of the engrafted sheep mandible. Accordingly, the material properties of the NiTi plates, modulated by the porous region, to be utilized in this procedure were assigned in a finite element model in ABAQUS (Dassault Systemes, Waltham, MA) of the fixation plates as visualized in a VSP environment, which initially was Materialise (Leuven, Belgium) Mimics. The example fixation plate consists of two components: (a) an outer solid “shell” with threaded, countersunk, solid screw-eyes for Ti-6A1-4V locking screws (provided by KLS Martin) and (b) an internal porous “core”, where the pore geometry is used to modulate stiffness. The pore geometry utilized orthogonally oriented struts of 300 mm diameter interspersed by 300 mm pores. FIG. 7B shows an example of the fixation plate of the study with exemplary dimensions. An interactive finite element model (FEM) design process that limited micromotion at the osteotomy site to 50 mm led to a form-fitting plate that began posterosuperiorly, posterior to the graft, immediately anterior to the first premolar roots, that curves down to the inferior border of the mandible immediately anterior to the graft.Attorney Docket No. 103361-579WO1 T2024-046

[0107] The final plate design used in this study places three screws in the host mandible anterior and posterior to the bone graft and two in the bone graft. The plate is 3 mm thick, which is near the maximum thickness for current Ti-6A1-4V plates. Fixation plate design included a screw planning finite element model (FEM), as shown in FIG. 8. This FEM ensured that all screws were bicortical (i.e., had maximum capture of the proximal and distal mandibular cortices). See Table 1, below, showing the lengths of screws for each sheep and implant. The fixation plate is 6 cm long, whereas the bone graft is 1.75 cm. The “wrap-around” geometry of the plate helps ensure that loading will bring about compression, rather than tension or spreading, at the osteotomy sites. Sixteen plates were 3D printed for this study at the University of Toledo on a Phenix Systems PXM (3D Systems, Rock Hill, SC) Laser Powder Bed Fusion 3D printer using Ni 50.1 -Ti Eckart TLS GmbH (Bitterfeld, Germany [Altana AG, Wesel, Germany]) powder sieved to produce particle sizes from 20 to 75 mm. The fabricated parts were polished, and thread milling was performed on a 5-axis CNC.

[0108] Following the fixation plate design, a cutting guide for the 17.5 mm osteotomized bone flap and a NiTi bone graft fixation plate. A cutting guide was designed and 3D printed based on the desired osteotomy sites to ensure the cutting of the bone and placement of the screws at the intended locations. Thus, the sterilized cutting guide is placed on the exposed bone, and surgical markings are made for screw tapping and osteotomies prior to placement of the graft fixation device.

[0109] Prior to opening the large animal model study, the virtual biomechanical parameters were verified via fatigue testing. A series of fatigue tests were conducted with cadaveric sheep mandible specimens using an MTS Systems (Eden Prairie, MN) Criterion Model 43 electric- controlled mechanical testing frame. A custom-designed rectangular fixture was designed and fabricated using 6061-T6 aluminum bars, steel fasteners, 304 stainless-steel threaded clamping connectors, and force applicators. The fixture featured an array of 40 holes on both the top and bottom plate that allowed for the placement of ’A” diameter threaded rods to surround the mandible sample in an orientation where the fixation repair location was facing up. The force applicator was then used from above to apply cyclic loading of 20N to the bone surface located just proximal to the fixation plate location at a rate of 2 Hz for 200,000 cycles. No significant separation or shifting of the bone relative to the plate was observed after testing for nearly 28 hours. This test was repeated three times for a NiTi fixation plate and a Ti-6A1-4V fixation plate. No significant shifting was observed in any of the test cases. Finally, the cleaning protocol was validated with cytotoxicity testing compliant with the ISO 10993-5 standard using L929 cells.Attorney Docket No. 103361-579WO1T2024-046Table 1. Screw lengths for installed devices in sheep.Sheep Model: Surgical Protocol

[0110] Five mandibular graft fixation (i.e., 4 Personalized NiTi fixation plates versus 1 standard-of-care, manually bent, Ti-6A1-4V fixation plates) implantation surgeries were performed by the same surgeon (HE). Briefly, the procedure began after a sheep was anesthetized and intubated. Next, an orogastric tube was placed as a vent to control for eructation. Sterile technique was followed beginning with the skin and fascia being incised to access the diastema region of the mandible. Using the aforementioned, personalized cutting guide, two lines were marked on the bone to guide the positioning of the fixation plate screws, and bone-cutting locations and screw holes were marked and then tapped. Next, the fixation plate was screwed in place. This was followed by the use of an oscillating saw to cut the two osteotomies previously marked 17.5 mm apart in the diastema region of the left mandible. These two osteotomies completely freed the previously marked bone graft. Since the cutting is done with the screws and plate in place, the gap between the bone and graft is uniform. The osteotomies were performed with careful attention to the inferior alveolar neurovascular bundle. After ensuring that there were no blood vessels injured; the periosteum, fascia, and skin were then sutured and closed in layers. The oral cavity was inspected for any unintended puncture wounds to the oral mucosa. None were observed in all 5 surgeries. In all 5 cases, this procedure lasted less than 4 hours. Along with post-operative care and monitoring, the oral cavity was flushed daily with an antiseptic solution for at least a couple of days. The sheep were assessed regularly, several times per day for 3 days, and at least once daily thereafter, for any complications and for wound and bone healing. Given that most large mandibular fixation plates fail from 1 to 3 years after implantation, the decision was made to allow animals 3 and 4 to survive until 1 year.Attorney Docket No. 103361-579WO1 T2024-046

[0111] Tracking Post-operative Bone Healing - Graft healing was tracked using histological and 3D CT. Post-operative 3D CT scans were taken within 3 months following surgery and periodically thereafter, as often as every 2 weeks, until complete healing was achieved. Fixation plate integrity, callus formation and retreat, and osteotomy bridging were carefully tracked via 3D CT.

[0112] Sacrifice and Histology - Immediately following sacrifice, the engrafted mandible was dissected and preserved in formalin. Prior to shipping to the Orthopaedic Research Bioengineering Laboratory at Colorado State University, each mandible was cut in half. One control, an unoperated mandible was processed, along with 4 experimental and 1 control hemimandible with standard-of-care Ti-6A1-4V graft fixation. Prior to histological processing, all hemimandibles were micro-CT-scanned. Hard plastic embedding was performed. Matching sagittal micro-CT and Sanderson’s Rapid Bone Stain and counterstain with Van Gieson stained sections were prepared.Results and Discussion

[0113] While all four experimental animals demonstrated complete mandibular graft healing, there were advances in technique between each of the four animals which led to faster healing, lack of screw loosening, and reduced callus formation. The first and third animals showed screw loosening. Two screws loosened in the first sheep, which was suspected to be due to those screws not being bicortical. Screw planning technique constantly improved from experimental animal sheep 1 through sheep 4. There was no screw loosening in Sheep 4. The second animal was smaller than the other 3. While the graft fixation plate fits well at surgery, the surgeon noted that it required more work than the other 3 to situation the plate. Sheep 2 was healing in the expected manner until, well into the healing period, it was observed that the fixation plate had broken into two pieces, each of which had moved away from the other. It was anticipated that this happened due to the remodeling of bone pushing away and breaking the more flexible fixation plate. Sheep 5’s Ti-6A1-4V fixation plate fractured rapidly, allowing a significant jaw dislocation. It was anticipated that this animal would need to be euthanized. However, it was able to chew well enough, likely on the unaffected side, to survive to the planned euthanasia.

[0114] All osteotomies of experimental sheep showed successful callus formation with bone remodeling to achieve osteotomy site bridging followed by full restoration of original bone anatomy and histology. Unexpected post-operative events were not observed. Complications in the form of infection, inflammation, or dehiscence were not observed. Pain and weight loss wereAttorney Docket No. 103361-579WO1 T2024-046 not out of the normal range. Sheep adversities were negligible, which included difficulty adjusting to the weather as their fur grew after each shear and some skin abscesses that healed rapidly.

[0115] The overall observation is that a well-placed, stiffness-matched plate brings several benefits. For example, reduced stiffness allows better engagement of the whole plate, avoiding stress concentrations. Additionally, better engagement of the whole plate and accurate screw planning reduce screw loosening is provided. Instead of a screw that is adjacent to a heavily loaded screw being loosened, it is tightened. This tightening is enhanced through accurate bicortical and changing screw orientation (i.e., not all in the same plane) when the plate wraps around the bone. Carefully allowing the plate to wrap around the bone helps keep the host and graft bone fragments in compression during loading, thus speeding the healing process.

[0116] Manually bent, off-the-shelf fixation plates usually have limited crimping areas that are located between the screw eyes. This compromises the elasticity of the plate in these areas, especially if repeated bending to reach the required form-fitting shape causes a work hardening at the most bent sites. Work-hardened sites may be prone to fatigue (i.e., caused by cyclic loading) failure. A flush fixation plate, with no gaps between it and the underlying surface, can act more as a unit than a plate with stand-off areas.

[0117] It is anticipated that a biomechanically-informed VSP session that simulates postoperative performance can optimize fixation plate location, shape (i.e., internal pore geometry and external surface shape), and material. In addition to providing cutting guides or intraoperative guidance to deploy personalized fixation, the VSP can help minimize the post-healing interruption of normal stress-strain trajectories. Allowing restoration of normal loading may avoid future stress shielding-induced bone loss or stress-concentration-induced device failure, especially cyclic loading-induced failure.

[0118] Conclusion - A novel stiffness-matching approach is presented herein. Rather than merely reducing stiffness by choosing NiTi over Ti-6A1-4V or matching stiffness to the adjacent bone, this approach matches the stiffness of the construct to the work of maintaining compression between adjacent healing bone segments. That work is also facilitated by engaging the full fixation plate / screw / bone fragment complex in order to reduce osteotomy site micromotion and speed bone bridging. That this strategy has now proven successful in a large animal mandibular graft model indicates to us that the approach shows promise for clinical mandibular graft and skeletal reconstruction applications that deserve future study.Attorney Docket No. 103361-579WO1T2024-046EXAMPLESThe Role of Stiffness-Matching in Avoiding Stress Shielding-Induced Bone Loss and StressConcentration-Induced Skeletal Reconstruction Device Failure

[0119] It is well documented that overly stiff skeletal replacement and fixation devices may fail and require revision surgery. Recent attempts to better support healing and sustain healed bone have looked at stiffness-matching of these devices to the desired role of limiting the stress on fractured or engrafted bone to compressive loads and, after the reconstructed bone has healed, to ensure that reconstructive medical devices (implants) interrupt the normal loading pattern as little as possible. The mechanical performance of these devices can be optimized by adjusting their location, integration / fastening, material(s), geometry (external and internal), and surface properties. This review highlights recent research that focuses on the optimal design of skeletal reconstruction devices to perform during and after healing as the mechanical regime changes. Previous studies have considered auxetic materials, homogeneous or gradient (i.e., adaptive) porosity, surface modification to enhance device / bone integration, and choosing the device’s attachment location to ensure good osteointegration and resilient load transduction. By combining some or all of these factors, device designers work hard to avoid problems brought about by unsustainable stress shielding or stress concentrations as a means of creating sustainable stress-strain relationships that best repair and sustain a surgically reconstructed skeletal site.

[0120] 1. Introduction - Metallic skeletal hardware is used for primarily two types of skeletal reconstructive surgery applications, bone fixation or bone replacement. Orthopaedic and craniofacial skeletal fixation or replacement surgeries are among the most frequently performed, and their rate of increase is among the highest of all surgical procedures. In 2017 approximately 22.3 million reconstructive orthopaedic surgeries were performed worldwide. The global orthopaedic device market size was valued at USD $40.9 billion in 2021 and is expected to expand at a compound annual growth rate (CAGR) of 3.1% from 2022 to 2030. Biometals are a crucial component in the development of medical implants and fixation devices. As such, it is imperative that they are composed of elements that are both nontoxic and nonallergic to the human body. In addition, biometals should possess high mechanical properties, wear resistance, and resistance to corrosion in order to prevent the dissolution of metallic elements. Finally, biometals should have a low stiffness that is similar to that of bone. These requirements ensure the safety and efficacy of biometals in the medical field. The principal metallic biomaterialsAttorney Docket No. 103361-579WO1 T2024-046 currently used in orthopaedic and craniofacial fixation or replacement devices consist of titanium alloys and, more rarely, stainless steel, nickel -titanium and cobalt alloys. Indeed, orthopaedic and craniofacial metallic devices play an important role in bone graft fixation for trauma, cancer, and arthritis reconstructive surgery.

[0121] 1.1 Background - While they have very different roles during the post-surgical healing period and the subsequent post-healing period, skeletal replacement and skeletal fixation devices may disrupt the normal stress-strain trajectories that bring about healthy, sustainable remodeling of cortical bone. Although standard-of-care skeletal devices ensure recovery in the majority of patients, the long-term prognosis for patients with significantly interrupted loading patterns includes a high risk of stress shielding-induced bone loss, stress-concentration-induced device failure (e.g., screw loosening, plate exposure, and screw or plate fracture), or both. High stress concentration on the hardware causes accelerated striping wear and fatigue, which might consequently result in its fracture or displacement. Meanwhile, stress concentration in the surrounding bone results in bone resorption or insufficiency fractures. Stress shielding is the reduction of bone density due to the redistribution of load that occurs when a skeletal device takes over the load on the bones. According to Wolff’s law, healthy cortical bone maintains its shape and radiodensity due to a continuous remodeling process. Healthy remodeling is evidenced by the ratio of osteoclasts to osteoblasts. That ratio is a response to normal mechanical loading and the continuous presence of micro-fracturing and its repair (i.e., remodeling). The skeletal device, constructed of a material that is much stiffer than the surrounding bone, supports major of the patient’s weight, thereby demonstrating an unphysiological redistribution of force transmission and preventing the local bone from receiving sufficient mechanical stimulus during remodeling. As a result, the bone will tend to redistribute and resorb. A decrease in bone mass and density may contribute to implant loosening and / or failure. The resorptive process generated by stress shielding has been reported as one of the main problems that lead to bone loss and revision surgery in craniofacial, knee, shoulder, spine, and hip implants, and bone fixation plates. Therefore, mechanical properties and geometry are critical factors that should be considered during skeletal device design to address stress shielding and stress concentration.

[0122] The concept of medical device stiffness-matching is not well-established in the medical or device engineering communities. Conceptually, it is important for both communities to consider the performance of a skeletal device and the bone it is attached to as a construct which is used under physiological loading conditions. It is best not to conflate the stiffness of aAttorney Docket No. 103361-579WO1 T2024-046 particular medical device or the reconstructed device and bone region with Young’s modulus of the test coupons of the dense (i.e., non-porous) material used in a fixation or bone replacement device. When describing the properties of isolated metal components, it is most common to use Young’s modulus. Thus, stiffness is a structural property, it is the resistance to elastic deformation, which is influenced by the geometry of the device / bone region as well as the materials of which it is comprised. From this perspective, the reconstructed region could be considered as a multi-material device that is affected by the way the load is applied. Young's modulus (or elastic modulus) is a mechanical property that measures a fully dense, material’s ability to withstand changes in length when under lengthwise tension or compression. It is intrinsic to a particular material and is not influenced by the geometry of a medical device, a reconstructed region, or how the load is applied to both. However, Young's modulus is, in essence, the stiffness of a device and is proportional to overall stiffness (i.e., the higher Young’s modulus, the higher the device’s stiffness). Bone has an anisotropic structure, which results in anisotropy, that is, mechanical properties depend on the direction in which they are measured. This can be explained using a relatively simple example of femoral cortical bone. The Young’s modulus of femoral cortical bone measured along the longitudinal direction is often approximately 18 GPa, while along the transverse direction, it varies between 5 and 10 GPa. However, in the case of bone, Young’s modulus is not as relevant as stiffness. Stiffness is a more representative factor that describes the mechanical properties of the bone since it incorporates the bone structure, geometry, and Young’s modulus. Many factors influence bone stiffness, such as age, sex, state of health, the relative amount of cortical and trabecular bone, normal loading pattern, the amount and direction of load experienced during behavior, and the anatomical site. The stiffness of different bones in the body varies from well below 3 GPa for some regions of trabecular bone, to often more than 20 GPa for the cortical region of large bones. Moreover, in the areas between hard and soft tissue — such as at the interface between cartilage, ligament, or tendon, and bone — the structure of adjacent soft tissue and bone can include a gradient of strength and elasticity. The Young’s modulus of the metals commonly used for skeletal repair or replacement devices, such as titanium alloys (e.g., Ti-6A1-4V), cobaltchromium alloys, and stainless steel are, approximately 116, 190, and 210 GPa, respectively, which makes them 4 — 10 times stiffer than the bone to which the device is attached. Although highly stiff skeletal devices may provide strong immobilization immediately after surgery, their long-term use may redistribute the load needed to sustain the repaired bone. This disruption in the bone’s normal stress-strain trajectories can lead to one or more of three poor outcomes: (a)Attorney Docket No. 103361-579WO1 T2024-046 stress shielding and bone resorption, (b) stress concentration in the fixation device and device failure (e.g., plate or rod cracking or screw pull-out), and / or (c) failure to restore muscle power as the reconstructed bone heals. After the bone is healed and strength returns to the muscles that allow it to operate, the possibility of these three risk factors increases if remnant fixation hardware continues to redirect bone loading significantly or if bone replacement hardware fails to re-establish sufficiently normal loading of the remaining bone. When the bone and / or reconstructive hardware fails, reoperation itself may lead to further hazardous consequences (e.g., more bone loss and pain at graft donor sites, scar tissue and poor vascularity, additional and more difficult reoperation often due to the same phenomenon caused by stiffness mismatch). Therefore, reduction, if not elimination, of the need to reoperate is a primary objective of personalized (i.e., in shape, location, and material), stiffness-matched skeletal reconstruction (i.e., fixation or replacement) devices.

[0123] 1.2 Current status and challenges - Bone function may be restored by partial or full replacement of diaphyseal (shaft) or epiphyseal (joint, e.g., hip, knee, shoulder, hand, foot, or spine) skeletal segments, clinically referred to as arthroplasty. The amount of motion that can be restored is a consideration. Science has allowed nearly full motion in hip, knee, and shoulder restoration, whereas intervertebral disc failure is usually treated by replacement (e.g., intervertebral fusion cage or scoliosis correction) with permanent skeletal immobilization instrumentation. Fracture and graft fixation can be problematic as surgeons are often unable to adequately anticipate or sufficiently personalize the device’s mechanical properties and performance. While device personalization is commonly available, often at great cost, clinicians rarely have access to mechanical information on the patient’s current or planned reconstructed skeletal anatomy or the loads it can be anticipated to encounter post-operatively. It would be useful for surgeons to have a way to anticipate the performance of a particular device irrespective of the level of personalization. Additionally, the performance must anticipate at least two mechanical goals for which the use of fixation hardware should not be in conflict as they occur at different points in the healing cycle. If planned well, good fixation can promote the transition from the initial demands of fixation during healing to the subsequent mass-sustaining bone loads that can lead to the long-term stability of the newly healed, reconstructed bone. Virtual elimination of motion at the osteotomy site is the job of fixation hardware. In the bestcase scenario, a fixation device will translate all load to the host and grafted bone as compression between these bone fragments. Minimization of torsion at the osteotomy site will allow and stimulate bone to quickly bridge the osteotomy site with minimal formation of callus.Attorney Docket No. 103361-579WO1 T2024-046 The completion of healing is also intended to restore the normal geometry and function of the newly healed bone both at the osteotomy site and throughout the affected anatomical segment(s). Maintaining compression at osteotomy or fracture sites will aid the healing of bone tissue and facilitate strengthening through remodeling. The planning and design of fixation must often account for healing and subsequently healed bone to be able to accept torsional load. Therefore, fixation hardware should be designed to prevent tension (i.e., twisting or splaying) at the osteotomy site under these and all loading conditions, including perhaps, a patient’s accidental fall. In designing fixation hardware and choosing its location, one would also look to allow the transfer of as much load encountered by the fixation plate, rod, or pin to its normal location in the bone once the bone has fully healed. It may be that the best design of a non-resorbable fixation device will be able to allow a sufficient load to pass normally, as if it were not there, through the healed bone for purposes of bone mass maintenance. Currently, physicians cannot attempt to design the location, shape, and material used in fixation devices to achieve the transfer of load following bone healing. Thus, the restoration of the normal stress-strain trajectory, while simultaneously minimizing host-graft bone osteotomy site micromotion at the early stages of bone healing, captures the overall challenge.

[0124] As noted, the choice of a fixation device’s external shape (pattern, length, width, thickness) and internal shape (porosity and pore geometry), location, fit to the bone, and material properties (materials used) of a skeletal replacement or fixation device may allow healing to occur but thereafter can lead to the mechanical failure of the device. A broken skeletal fixation or replacement device requires surgical removal and replacement with sufficient bone and fixation hardware to support healing and sustainable bone remodeling thereafter. A nonpersonalized approach is usually blind to stress concentration in the hardware and the potential for thread-stripping wear between the screws and the fixation plate, screws pulling out of the bone, or fixation plate fatigue failure (i.e., cracking due to cyclic loading) which might consequently result in the device’s fracture, displacement, and bone segment dislocation. In addition to device failure, failing to consider the mechanical performance of the fixation plate can lead to stress shielding of the surrounding bone. By redirecting the normal bone loading pattern, the bone that previously received a sufficient load to sustain morphology-preserving remodeling may resorb and, potentially, incur insufficiency fracture.

[0125] Currently, physicians have limited control over the personalization of the external shape of these devices and, thus, where they can be optimally located. Further, surgeons have virtually no input into the device’s material properties, such as their constituent material, theAttorney Docket No. 103361-579WO1 T2024-046 amount of porosity in that material, or the pore space’s geometry. If the device vendor chooses to construct the device from a biocompatible but highly stiff material, that device may serve the role of immobilization during healing, but during that time and after healing, it may redirect stress-strain trajectories in unphy siological directions that may give fixation hardware an unsustainable load and / or may prevent the healed and adjacent bone from receiving sufficient mechanical stimulus to bring about a healthy, sustainable remodeling process. Interrupting the normal stress-strain trajectories may also lead areas of thick, load-bearing cortical bone to resorb through remodeling as the load they previously received is redistributed elsewhere. When bone receives an insufficient load, it is stress-shielded. Stress shielding can lead to bone loss. Because the device cannot remodel (i.e., cannot repair itself), cyclic loading can lead to device fatigue failure or loosening of the device, which is associated with revision surgery for craniofacial, knee, shoulder, spine, hip implants, and bone fixation plates. Therefore, mechanical properties, geometry, and the location of a bone replacement or bone fixation device can be used to avoid the negative outcomes associated with bone stress shielding and device stress concentration. For example, see FIG. 9 showing a schematic illustration of optimization of the stiffness-matched orthopedic device and virtual surgical planning. Current manufacturers of personalized joint replacement and fixation devices rarely provide physicians with biomechanical data on these parameters for a device their patient will receive. Simulations of predicted device performance could provide physicians confidence in skeletal reconstruction device decision-making once information on outcome optimization (i.e., anticipated performance) was available.

[0126] Surgeons are generally aware that overly stiff skeletal replacement or fixation devices carry a high risk of failure. However, they may not be aware that there is a remedy for this widespread truism. There is often some confusion around concepts of stiffness-matching as it is often assumed that the best device, especially a fixation device, will have uniform strength and stiffness. For example, an oversimplified view would be to expect that the optimal condition always requires a device that has the same stiffness as the bone to which the device is attached. However, with fixation hardware, this is not optimal as the device is not replacing the bone. Rather, it may be best if it were somewhat stronger and stiffer than the surrounding bone, sufficiently so to minimize micromotion at the osteotomy (healing) site. Thus, an optimal stiffness-matched fixation concept is likely to conceive of hardware that is matched to the need to immobilize healing bone segments during loading by bringing them into compression. Allowing tension at the osteotomy (healing) site would be detrimental. Further, the optimal location for fixation is likely to accomplish this goal without, or with only minimally,Attorney Docket No. 103361-579WO1 T2024-046 interrupting the normal, bone-preserving, stress-strain trajectories that are prevalent after the bone has healed. Finally, it is expected that the optimal stiffness would be the minimum needed to accomplish the goals of fixation so that as much as possible of the whole plate-fixation-bone complex could be engaged in compression during loading and therefore, prevent stress concentration in the device and drive remodeling through compression of the healing bone. An overly stiff fixation plate would be more likely to concentrate load around a screw or a relatively weak point in the plate or rod, thereby increasing the risk of failure due to stress concentration and / or stress concentration plus cyclic loading (i.e., associated with fatigue failure). Material choice, careful screw planning (e.g., bicortical screws inserted at complementary, not uniform, angles), desirable surface modification, and strategic use of porosity can also help aid healing and prevent device failure.

[0127] One of the core aspects of the concept of stiffness-matched fixation devices is that they both facilitate the immobilization of reconstructed bone (e.g., skeletal fixation) and, following healing (i.e., fixated osteotomies heal), they facilitate the healed bone receiving a normal load (i.e., they do not re-direct that load, or only do so minimally). In the case of fixation, this means that the normal load is transferred to the repaired bone as it heals, often in association with the restoration of muscle power and mobility. Two principles are helpful here. First, it is recommended to start with devices no stiffer than they need to be to accomplish their function. In skeletal fixation, that stiffness must be sufficient to redirect all loading to keep the healing osteotomy sites in compression, not tension. Second, the device should be flexible enough to engage the entire construct during loading, as local stress concentrations are more likely to lead to device failure. External surface and internal pore geometries can also be used to ensure restorative loading patterns that promote healthy bone loading and sustainable remodeling.

[0128] Overall, one should envision stiffness-matched devices with personalized shapes and mechanical properties that restore the normal stress-strain trajectories of the repaired bone and match the stiffness to the work the device must perform in concert with the bone. Thus, the restorative device’s properties are tailored, and personalized to work in concert with the host’s anatomy and the long-term needs of the patient. Stiffness-matching can be achieved by different approaches: application of materials with modulus close to that of the bone, adaptation of geometry with pre-defined mechanical properties, and / or surface modification that ensures good integration and proper load transferring to the bone. For example, see FIG. 10 showing a diagram of the basic approaches to a stiffness-matching objective.Attorney Docket No. 103361-579WO1 T2024-046

[0129] 2. Low-modulus materials - The initial approach to skeletal reconstruction devices was driven by a failsafe approach to mechanical properties and careful attention to biocompatibility together with high corrosion resistance. Among metallic materials, titanium and its alloys have recently been considered the most suitable material for skeletal reconstruction devices, as they rapidly self-passivate and are less problematic than competing materials, such as stainless steel 316L, Cr-Co alloys, and tantalum. For example, see FIG. 11 showing a graph of modulus of elasticity of biomedical alloys. Further, Table 2 provides a comprehensive overview of the advantages, disadvantages, and challenges associated with the use of common biomaterials in orthopedic implant applications. More recently, low-modulus titanium alloys, with various alloying elements, such as Nb, Mo, Zr, Ta, Mn, and Cr, have been found to improve mechanical performance. Low modulus alloys used for skeletal implants include, but are not limited to, Ti- 13Nb-13Zr, Ti-12Mo-6Zr-2Fe, Ti-15Mo alloys, Ti-45Nb and Ti-35Nb-7Zr-5Ta. Extensive research has been performed on Gum Metal (TNTZ), which is a unique class of P-type titanium alloy. Initially, TNTZ was composed primarily of titanium, 23% niobium, 0.7% tantalum, 2% zirconium, and 1% oxygen. It can exist over a range of compositions, including those with vanadium and hafnium contributions. These alloys simultaneously offer high strength (> 1 GPa), a low Young’s modulus (~60 GPa), and good biocompatibility. Moreover, shape memory NiTi (aka Nitinol or nickel -titanium) alloys have found wide applications after initial problems with significant nickel ion release. Nevertheless, the release of harmful allergenic and cytotoxic ions from titanium alloys, stainless steel, and Co-Cr alloys has been previously reported as well. However, after improvements in alloying, NiTi has gained acceptance as a standard-of-care material for orthodontic arch wires, vascular stents, active catheters, trans-aortic valve replacement rings, and most recently, for skeletal applications such as spine-fracture staples, hallux restraint, and lumbar vertebral fixation components. Given the difficulty of forming NiTi, these devices have often been primarily formed via laser cutting or as drawn wire. However, more recently, the use of 3D-printed NiTi for bulk devices has been explored. Biomedical interest in NiTi alloys, especially for bone restoration applications, derives from the low Young’s modulus of these alloys (austenite: 70 GPa; martensite: 28-40 GPa) and their biocompatibility when Ni ion release is minimized through careful alloying and surface treatments. Similarly, biomedically relevant magnesium alloys present a low modulus of elasticity value (~40 GPa), one that is close to that of cortical bone (—15-30 GPa). Moreover, in addition to biocompatibility and desirable mechanical properties, biomedically-relevant magnesium alloys have the added benefit of being resorbable. Recently, there has been extensiveAttorney Docket No. 103361-579WO1 T2024-046 research into resorbable magnesium alloys that will have sufficient strength for use in skeletal fixation hardware and will not degrade before the surgically reconstructed bone has healed.Table 2. Most common biomaterials for orthopedic implants.

[0130] 3. Surface conditions - The successful integration of medical devices into the human body heavily relies on the properties of their surfaces. In order to enhance biocompatibility, as well as mechanical, tribological, and antibacterial properties, a range of modifications can be applied to the surface. It is important to note that surface roughness at different scales, including macro, micro, and nanoscale, can have a significant impact on cell behavior. Studies have shownAttorney Docket No. 103361-579WO1 T2024-046 that osteoblasts tend to adhere and spread more effectively on surfaces that have micrometric roughness. The enhancement of the implant surface has been shown to improve the initial process of osseointegration with the alveolar bone, resulting in a reduction in the duration of treatment. Osseointegration is a critical process in which newly formed bone tissue adheres to the surface of a bone implant. The rate at which the initial process, as well as ongoing remodeling, occurs must be sufficient to ensure that the implant remains stable within the body. The level of remodeling will likely depend on the load the attachment site receives over time. This phenomenon has been observed since the first use of bone implants in medical procedures and is of utmost importance in orthopedic applications. Furthermore, alterations to the surface of implants that result in an increase in surface area for attachment to bone have been found to enhance bone-to-implant contact (BIC). Techniques that can be employed to achieve this include increasing surface roughness or surface porosity (i.e., textured surface). Surface porosity can be generated utilizing methods such as acid etching or anodizing. Electron Beam Powder Bed Fusion is also recognized as a means of creating surface porosity. When surface porosity is established at the nanoscale and microscale level, it can significantly improve the absorption of biological fluids immediately following implantation. This, in turn, accelerates the growth of BIC dynamics, leading to quicker loading of the implant and an enhanced maximum value. Ultimately, this facilitates the implant's ability to bear larger loads once the osseointegration process is fully developed.

[0131] The stability of a replacement implant’s integration or a fixation device’s attachment is a critical factor in determining its success. This stability is assessed at two stages: primary and secondary. Primary stability is achieved right after surgery (i.e., prior to any healing, new bone formation, or bone strengthening) and relies on the quality of the bone at surgery and the interaction between the bone and implant at the bone-implant interphase. On the other hand, secondary stability is attained during the healing period through the osseointegration process. This process involves the creation of new bone that tightly bonds with, if not infiltrates, the implant and matures over time. A 2-D finite element model has been developed for the boneimplant interphase to evaluate the impact of the bone Young’s modulus, implant surface roughness, and the implant’s material composition. For example, see FIG. 12 showing schematic illustrations of the 2-D finite element model and spatial variation of the maximal shear stress in the periprosthetic bone tissue for different implant roughness and bone-implant contact ratios. Lowering the BIC ratio leads to an increase in stress shielding, which can negatively impact the clinical outcomes of endosseous implants. Thus, it is imperative to model and then prioritizeAttorney Docket No. 103361-579WO1 T2024-046 maximizing providing a sufficient BIC ratio to avoid stress shielding. Moreover, it was shown that matching the mechanical properties of bone tissue with implant material properties and morphology can result in a uniform shear stress field, which may aid in avoiding stress-shielding effects.

[0132] As already noted, in order to promote healing, both skeletal replacement and skeletal fixation devices must become integrated, literally fastened, with the bone to restore normal stress-strain trajectories. Their integration will determine the reconstructed region’s ability to efficiently transduce load between surrounding muscles, ligaments, host bone, grafted bone (if any), screws, and a fixation device. There are many ways that “fastening” can be accomplished. However, the most common are screws, a press-fit shape for a complex surface, bone “cement”, and / or bone growth into the device. Indeed, more than one type of fastening process can be used at various locations of an implanted skeletal repair device. As noted, osseointegration of skeletal replacement devices may benefit from surface texturing and / or coatings. Surface modification may include but is not limited to coatings, layers, or surface texture (i.e., roughness). While rough surfaces may promote bone ingrowth (osteoconduction), they may also promote bone ingrowth into a porous region (i.e., osteoinduction). Metallic skeletal device coating techniques include physical vapor deposition, chemical vapor deposition, electrochemical deposition, solgel coatings, plasma spraying, and micro-arc oxidation. In addition to creating rough surfaces, ceramic or polymer coatings may contain bone or vascular growth factors, immune-modulating factors, or antibiotics. However, mechanical failure is to be avoided; in all cases, sufficient implant surface integration is necessary at locations where load transfer occurs. High levels of micromotion should be prevented at both the bone-implant interface, screw / device interface, and at all osteotomy sites where the bone is expected to heal. Micromotion above 300-400 pm between the device and adjacent host bone and especially at the healing site between two osteotomized bone fragments may prevent bone bridging and healing its closure (i.e., nonunion). Integration will improve load transduction between adjacent bone and the host. In doing so, it will reduce any callus that has formed. However, bone healing may have no effect on stress shielding-induced bone loss or stress concentration-induced device failure. The importance of maximizing bone-implant contact surface area has been shown, especially relative to areas where the load is transduced between the host bone and a skeletal fixation or replacement device. Careful attention to modelling and achieving anticipated loads at this interface will facilitate the device / bone construct’s integration and thus, its role in carrying a load in a skeletalAttorney Docket No. 103361-579WO1 T2024-046 replacement device or, in the case of a skeletal fixation device, maintaining osteotomy site compression and avoiding osteotomy site motion that would prevent healing.

[0133] 4. Geometry - Geometry pertains to the overall shape of the medical device, including its external and internal features. The design of a medical device is determined by a multitude of factors, including the space it must occupy (in the case of implants) or the surface it must be attached to (in the case of fixation devices). In the development of medical devices, achieving targeted stiffness levels necessitates careful consideration of the device's shape and mechanical properties. Such requirements have a significant impact on the presence or absence of internal features, including truss structures and porosity, within the device's design.

[0134] 4.1 Defined internal porous geometry and additive manufacturing - At best ordered, but usually random, porosity with varying permeability to bone or body fluids is the best outcome from the inclusion of porogens or other poorly controlled techniques for producing porous space (i.e., initially air and later, inside the body, fluid-filled space) in metallic medical implants. Somewhat ordered but still primarily random, pore geometries can be fabricated via conventional methods such as powder metallurgy, chemical or electrochemical deposition, vapor deposition, and liquid-phase methods. As a class, these regions are often referred to as foam structures. The inclusion of these pore spaces will reduce the stiffness of the device, but it is difficult to predict their final mechanical properties given the inhomogeneous and relatively unpredictable formation of porosity. However, the development of CAD / CAM technologies, especially additive manufacturing, has allowed the fabrication of devices with pre-defined isotropic and anisotropic internal pore geometry and surface porosity (i.e., defined texture). This has created a revolution in stiffness-matching (see Section 4.3). There are two commonly used approaches to designing porous structures. One involves direct (i.e., personalized) design through the use of patient images, while the other involves indirect design through the use of parametric porous units. Defined internal porous geometry incorporates precise specifications for pore morphology, such as strut shape and tortuosity, strut size (e.g., diameter) and distribution pore shape and tortuosity, as well as pore size and pore size distribution. Moving from simple lattices where strut junctions are high angle stress concentrations (e.g., orthogonal) to octahedral and Schwartz “D” (diamond) repeating units, aka “pore cells”, through a triply periodic minimal surface (TPMS) geometries such as Schoen’s gyroid, Voronoi-Tessellation, and other auxetic structures. Current research focuses on anisotropic pore geometries, which cannot only be used for stiffness-matching but also to direct the load. Numerical analysis allows one to estimate mechanical performance in situ as part of a Virtual Surgical Plan (VSP) as wellAttorney Docket No. 103361-579WO1 T2024-046 as the permeability of a device. In situ mechanical properties of bone can be estimated from the geometry and inferred material constituents of musculoskeletal structures (i.e., bone, cartilage, muscle, tendon, and ligament) seen in 3D MRI, 3D ultrasound, or, more commonly, 3D CT images. Moreover, recent studies have revealed that porous scaffolds with lower stiffness in comparison to the bone they are intended to replace typically lead to a greater level of bone ingrowth. For example, see FIG. 13, with panels A, B, and C, showing A. CT-Renderings of bone ingrowth into a locally stiffness-matched porous scaffold (not shown); B. Fluorescent Microscopy Images for animals 4 and 6 with islets of bone deep within the scaffold interior circled; and C. Histology Images for animals 4 and 6 where A and B are magnified sections of the scaffold periphery and interior respectively. Examples of osteocytes are circled.

[0135] Surprisingly, mechanical personalization is often touted. However, physicians are rarely shown data on the anticipated mechanical performance of a device in their patient in an interactive VSP environment. Rather, consideration of a personalized device’s mechanical performance in a particular patient is left to a confidential, company-specific, often trade secret and regulatory agency-approved device design envelope that may not include any personalization but rather the anticipation that it will be sufficient for a wide range of mechanical conditions. Often, this certainty is gained, indeed claimed in an application for commercial clearance, from validation studies against known standards (e.g., ASMT F382, ISO 9585). Thus, the extent of the personalization and optimization of mechanical performance, if any, is not known outside the company. This is likely to change as physicians gain the ability to model a fixation or replacement device’s location, shape, and material properties and perhaps manufacture the modelled device, at the Point-of-Care. Point-of-Care Manufacturing (POCM) does occur now at major medical centers (e.g., Walter Reed National Military Medical Hospital, MD, USA, and the Mayo Clinic, MN, USA, are among this select group). Once the physician- informed design and POCM become widespread, the medical literature will begin to accumulate device-specific data and the evidence available to physicians on which to judge predicted, personalized, and optimized functional outcomes, they will then be able to advise their patients with a personally validated recommendation, versus an experience or literature-based evidence, for a personalized skeletal reconstructive device design. As with the base metal’ s(s’) material properties, pore and strut size and geometry (i.e., shape) can be adjusted to personalize a device’s mechanical properties. Pore geometry need not be isotropic. It can be adaptive, for example, in order to keep healing bones in compression while subsequently transferring the load to the healed bone once the healing process is complete. Perhaps most relevant are previousAttorney Docket No. 103361-579WO1 T2024-046 attempts to reduce device stiffness through the strategic use of porosity, such as the use of additively manufactured square pore structures in a CoCrMo device that achieved stiffness and strength characteristics similar to cortical and trabecular bone in the human femur or the incorporation of less stiff materials inside a metallic shell. However, while theoretically, stiffness can be reduced by changing materials and / or geometry, it might be difficult to avoid deleterious stress concentration with high-angle strut intersections or very thin struts that may fail due to fatigue in the future. This could occur due to cyclic loading if sufficient load is not transferred to newly healed bone.

[0136] 4.2 Meta-biomaterials - The term ‘meta-materials’ refers to engineered materials exhibiting novel properties not usually found in nature. Biocompatible meta-materials with novel properties such as negative compressibility, negative Poisson’s ratio, and negative elasticity offer new opportunities. Auxetic structures (aka NRP- Negative Poisson’s Ratio) are a specific class of meta-materials, that have a negative Poisson’s ratio thanks to their extraordinary ability to laterally expand in response to axial tension. The medical variant of meta-materials is meta-biomaterials. These materials may have specialized applications in skeletal replacement devices if it is useful for the device to expand under pressure. Such expansion may allow the compression of adjacent bone by maintaining firm contact. For example, see FIG. 14, showing hybrid meta-biomaterials with (A) Test set-ups and image processing: (1) off-axis compression, (2) off-axis compression with an adherent layer used to improve visualization of strain distribution, (3) points plotted along the borders of a hybrid meta-biomaterial and the order in which they were numbered according to their position; (B) Horizontal strains in the tape surrounding hybrid meta-biomaterials types 1-6 at 2 mm displacement; and (C) Mean maximum expansion during off-axis compression. Extensive evaluations of auxetic meta-biomaterials have been made for skeletal reconstruction. For example, it was found that rationally designed metabiomaterials present an opportunity to produce a hip implant incorporating auxetic (negative Poisson’s ratio) and conventional (positive Poisson’s ratio) properties in order to maintain implant-bone contact and potentially reduce stress shielding-induced, so-called “aseptic” loosening.

[0137] The interface between the implant and bone is more likely to fail under tension compared to compression, based on described failure criterion. Additionally, bone is more resistant to compression than tension. If the implant’s Poisson’s ratio remains constant, one side will retract from the bone while the other is compressed against it. Considering existing criterion, the side that experiences tension (and therefore retracts from the bone) is more likelyAttorney Docket No. 103361-579WO1 T2024-046 to experience interface failure due to the different mechanical strengths of bone in tension and compression. Moreover, to prevent wear particles from entering the enclosed cavity, it’s crucial to ensure maximum contact between the implant and bone. Additionally, the fixation of the implant can be enhanced through osseointegration prompted by mechanical stimulation in accordance with Wolff’s law. Finally, studies have concluded that all of these points emphasize the necessity of devising an implant that exerts compression on both sides of its neutral axis. Femoral implants with regions of negative, positive, and graded Poisson’s ratio unit cells were studied. It has been concluded that materials with negative and graded Poisson’s ratios could be used to create a uniform distribution of micromotion at a bone-implant interface. A titanium auxetic meta-biomaterial has been designed and experimentally validated consisting of rotating cuboids connected by an optimized micro- structure, with bone-mimicking elastic modulus and low density for replacement of the region of trabecular bone in vertebral implants. Auxetic structures were also found to exhibit more effective stress transfer and attenuation under practical loading conditions compared with natural intervertebral discs and conventional 3D implants. Given that meta-biomaterials are being considered primarily for permanent skeletal replacement, it will be important to confirm the number of cycles that these properties can be maintained without fatigue failure. Similarly, if they are to be used in fixation devices, it will be important to ensure that they allow sufficient load transfer to the newly healed bone to avoid stress shielding or the creation of unsustainable stress-strain trajectories (i.e., maladapted for the bone, especially its cortical component).

[0138] 4.3 Tailored and functional gradient porosity - Functionally graded materials - In most cases where 3D-printed pore geometries have been explored, they have relied on uniform pore geometry (i.e., isotropic and uniform strut and pore diameters). Since most bones have preferred loading patterns, these patterns are reinforced by irregular and gradient distribution of cortical bone osteon structure. The number and density of such osteons will depend on local mechanical demands. Therefore, functionally graded pore geometry and / or materials may be able to mimic the hierarchical and gradient structure found in natural bone. This is particularly important for the long-term viability of skeletal replacement devices, where a model of the mechanical properties and predominant loading patterns of bone that is to be replaced by an implant could help one optimize the design of a skeletal replacement device. Moreover, graded structures could potentially improve osteointegration and wear resistance. Gradient structures can be obtained by a compositional, microstructural, or geometric gradient from one surface of the material to the other, resulting in a material with continuously varying properties that couldAttorney Docket No. 103361-579WO1 T2024-046 channel load in desired directions. Indeed, the use of porosity to create functional gradients by changing the size of the pores, trabeculae, or both of these features by uniform and smooth gradients has been extensively explored. The tailoring of porosity gradients in porous hip stems has been shown to be effective in reducing Young’s modulus and therefore, mitigating bone resorption secondary to stress shielding. For example, see FIG. 15 showing calculation and modelling of gradient porosity to yield desired mechanical properties for bone implants. Graded structures have also been implemented in dental implants to improve osseointegration.

[0139] 4.4 Design of stiffness-matching skeletal repair devices - Numerical analysis is the first step to determining the optimal location, shape (i.e., external surface and internal pore geometry, if any), and single or multi-material composition of a skeletal repair device. Choices in any of these parameters will interact. Additionally, the reconstructed optimization of these parameters should be guided by the performance of the eventual part in situ. Loading patterns (direction, frequency, and magnitude of load) must be accurately and sufficiently predicted to determine the risk of stress shielding or stress concentration. Moreover, the device + bone construct may need to include muscles, tendons, ligaments, and screws.

[0140] The 3D source data for this biomechanical scene may be used to suggest where both uniform and gradient porous structures may be used to direct sustainable loading patterns to promote healing and / or the integration of a skeletal replacement device such as total joint replacements at the proximal femur or tibial plateau. For example, porous hip femoral stems have been demonstrated effective in reducing bone resorption secondary to stress shielding. Ti- 6A1-4V segmental bone defect replacement implants have been modeled, and it has been found that unsheathed cellular designs can release metal debris. An implant that combines an anatomically matched geometry and a stiffness-matching strategy that would restore the original bone’s stress-strain trajectories has also been proposed. Low-stiffness NiTi alloys with engineered porosity have been studied to optimize personalized fixation plates with stiffness- matched regions that are attached to the adjacent bone. Stiffness reduction can be achieved through the use of 3D Topological Optimization. A stemmed tibial component has been demonstrated with a fully porous stem with a tuned tetrahedron lattice architecture. The porosity has been optimally tailored to mitigate postoperative bone resorption and end-of-stem pain while satisfying the strength requirement necessary to sustain cyclic loading. The numerical results suggest that the overall amount of bone resorption around a graded porous tibial stem is 26% lower than that around a conventional, commercially available, fully dense titanium implant of identical shape and size. A similar result was found in a multimaterial (i.e., thin metal sleeveAttorney Docket No. 103361-579WO1 T2024-046 around the plastic core) total hip replacement device. An interesting approach to restoring the normal loading of fixated bone following the healing of bone segments is the bone bandaid concept. There, the releasable fixation mechanism effectively reduces the fixation device’s stiffness, thereby preventing it from accepting a load after the bone is healed. The outcome here is similar to what would occur with a resorbable device. However, large, fully resorbable skeletal fixation devices must ensure that degrading fragments do not break free and lodge in adjacent tissues or joint cavities. The selected stiffness-matching skeletal repair devices are listed below in Table 3.Table 3. Overview of stiffness-matching skeletal repair devices

[0141] 5. Conclusions - As mentioned, the primary dichotomy between the performance of skeletal restoration devices is between skeletal fixation and skeletal replacement roles. While a device with a stiffness similar to that of adjacent bone will efficiently transduce load across the device-bone interface, this is not a sufficient design criterion to meet all of the demands of eitherAttorney Docket No. 103361-579WO1T2024-046 type of skeletal repair device. As mentioned, skeletal fixation must be located where it can fix initially freely moving bone segments in compression under any load, especially the highest experienced load, throughout the healing period. Thereafter, the device must minimize, if not eliminate, interference with the normal stress-strain trajectories that the bone’s shape and mechanical properties, as well as the adjacent muscles, are adapted to. Skeletal replacement devices must perform in this manner, i.e., to facilitate force transduction across the host boneimplant attachment site, as well as to ensure that the normal stress-strain trajectories of the remaining bone are not interrupted. Additionally, a skeletal replacement device must not undergo cyclic (fatigue) failure during its expected use life.

[0142] Recent research focuses both on the modelling of device location, shape, and materials and on finding new ways to manufacture devices that will bring about an optimal restoration of skeletal performance while reducing the risk of reoperation due to device and / or skeletal failure. In an effort to eliminate these problems, a holistic stiffness-matching approach to design and fabrication may be a promising option. As presented in this review, modification of the properties of the device can be achieved in several ways, such as modeling the mechanical performance of the relevant bone-device construct in a way that takes into account the forces it will undergo in the design of restorative devices, use of the materials with modulus close to that of bone, surface modification to enhance device / bone integration, auxetic materials to engender desired mechanical performance, homogeneous or gradient (i.e., adaptive) porosity for the same purpose, and ultimately by combining all or some of these factors to avoid problems brought about by unsustainable stress shielding and / or stress concentrations. During healing, it is critical that skeletal fixation minimizes micromotion between healing bone segments. In the posthealing period following skeletal fixation, or on implantation of a skeletal replacement device, it is critical to restore, not interrupt, normal stress-strain trajectories. In addition to novel surgical planning software that includes device design based on performance modelling, further research on materials that can change their properties over time or can become mechanically irrelevant after the reconstructed bone has healed, including partially inert and partially resorbable multimaterial devices (e.g. inert metal-resorbable metal, inert metal-resorbable polymer, inert metal- resorbable, metal composite, etc.) is needed.Configuration of Certain Implementations

[0143] The construction and arrangement of the systems and methods as shown in the various implementations are illustrative only. Although only a few implementations have beenAttorney Docket No. 103361-579WO1 T2024-046 described in detail in this disclosure, many modifications are possible (e.g., variations in sizes, dimensions, structures, shapes, and proportions of the various elements, values of parameters, mounting arrangements, use of materials, colors, orientations, etc.). For example, the position of elements may be reversed or otherwise varied, and the nature or number of discrete elements or positions may be altered or varied. Accordingly, all such modifications are intended to be included within the scope of the present disclosure. The order or sequence of any process or method steps may be varied or re-sequenced according to alternative implementations. Other substitutions, modifications, changes, and omissions may be made in the design, operating conditions, and arrangement of the implementations without departing from the scope of the present disclosure.

[0144] The present disclosure contemplates methods, systems, and program products on any machine-readable media for accomplishing various operations. The implementations of the present disclosure may be implemented using existing computer processors, or by a special purpose computer processor for an appropriate system, incorporated for this or another purpose, or by a hardwired system. Implementations within the scope of the present disclosure include program products including machine-readable media for carrying or having machine-executable instructions or data structures stored thereon. Such machine-readable media can be any available media that can be accessed by a general purpose or special purpose computer or other machine with a processor. By way of example, such machine-readable media can comprise RAM, ROM, EPROM, EEPROM, CD-ROM or other optical disk storage, magnetic disk storage or other magnetic storage devices, or any other medium which can be used to carry or store desired program code in the form of machine-executable instructions or data structures, and which can be accessed by a general purpose or special purpose computer or other machine with a processor.

[0145] When information is transferred or provided over a network or another communications connection (either hardwired, wireless, or a combination of hardwired or wireless) to a machine, the machine properly views the connection as a machine-readable medium. Thus, any such connection is properly termed a machine-readable medium. Combinations of the above are also included within the scope of machine-readable media. Machine-executable instructions include, for example, instructions and data which cause a general-purpose computer, special purpose computer, or special purpose processing machines to perform a certain function or group of functions.Attorney Docket No. 103361-579WO1 T2024-046

[0146] Although the figures show a specific order of method steps, the order of the steps may differ from what is depicted. Also, two or more steps may be performed concurrently or with partial concurrence. Such variation will depend on the software and hardware systems chosen and on designer choice. All such variations are within the scope of the disclosure. Likewise, software implementations could be accomplished with standard programming techniques with rule-based logic and other logic to accomplish the various connection steps, processing steps, comparison steps and decision steps.

[0147] It is to be understood that the methods and systems are not limited to specific synthetic methods, specific components, or to particular compositions. It is also to be understood that the terminology used herein is for the purpose of describing particular implementations only and is not intended to be limiting.

[0148] As used in the specification and the appended claims, the singular forms “a,” “an” and “the” include plural referents unless the context clearly dictates otherwise. Ranges may be expressed herein as from “about” one particular value, and / or to “about” another particular value. When such a range is expressed, another implementation includes from the one particular value and / or to the other particular value. Similarly, when values are expressed as approximations, by use of the antecedent “about,” it will be understood that the particular value forms another implementation. It will be further understood that the endpoints of each of the ranges are significant both in relation to the other endpoint, and independently of the other endpoint.

[0149] “Optional” or “optionally” means that the subsequently described event or circumstance may or may not occur, and that the description includes instances where said event or circumstance occurs and instances where it does not. Throughout the description and claims of this specification, the word “comprise” and variations of the word, such as “comprising” and “comprises,” means “including but not limited to,” and is not intended to exclude, for example, other additives, components, integers or steps. “Exemplary” means “an example of’ and is not intended to convey an indication of a preferred or ideal implementation. “Such as” is not used in a restrictive sense, but for explanatory purposes.

[0150] Disclosed are components that can be used to perform the disclosed methods and systems. These and other components are disclosed herein, and it is understood that when combinations, subsets, interactions, groups, etc. of these components are disclosed that while specific reference of each various individual and collective combinations and permutation ofAttorney Docket No. 103361-579WO1 T2024-046 these may not be explicitly disclosed, each is specifically contemplated and described herein, for all methods and systems. This applies to all aspects of this application including, but not limited to, steps in disclosed methods. Thus, if there are a variety of additional steps that can be performed it is understood that each of these additional steps can be performed with any specific implementation or combination of implementations of the disclosed methods.

Claims

Attorney Docket No. 103361-579WO1T2024-046WHAT IS CLAIMED IS:

1. A bone fixation plate comprising: a first end and a second end opposite and spaced apart from the first end along a longitudinal axis; a first surface extending from the first end to the second end; a second surface opposite from the first surface; a network of interstices disposed between the first surface and the second surface; and a plurality of screw holes defined by and extending through the bone fixation plate from the first surface to the second surface, wherein the network of interstices is configured to reduce an overall stiffness of the bone fixation plate as compared to a solid reference plate, the overall stiffness substantially matching that of an underlying bone on which the bone fixation plate is implanted.

2. The bone fixation plate of claim 1, wherein one or more characteristics of the network of interstices is selected based on a virtual surgical model of the underlying bone and adjacent anatomy of a patient.

3. The bone fixation plate of any of claims 1-2, wherein the network of interstices comprises a network of pores.

4. The bone fixation plate of any of claims 1-3, wherein each interstice in the network of interstices has either (i) a uniform shape and size or (ii) a non-uniform shape and size.

5. The bone fixation plate of any of claims 1-4, further comprising a plurality of struts extending between at least a portion of the bone fixation plate between the first surface and the second surface, the plurality of struts defining the network of interstices.

6. The bone fixation plate of claim 5, wherein the plurality of struts comprises a first plurality of struts extending longitudinally along the bone fixation plate and a second plurality of struts extending substantially perpendicularly to the first plurality of struts.

7. The bone fixation plate of any of claims 1-6, wherein each of the interstices in the network of interstices has a substantially rectangular shape.Attorney Docket No. 103361-579WO1T2024-0468. The bone fixation plate of any of claims 1-7, wherein each of the interstices in the network of interstices is substantially circular.

9. The bone fixation plate of any of claims 1-8, wherein the longitudinal axis of the bone fixation plate is curvilinear, and the first surface and the second surface have a curvature matching that of the longitudinal axis.

10. The bone fixation plate of any of claims 1-9, wherein the curvature of the first surface and the second surface matches that of an underlying bone on which the bone fixation plate is implanted.

11. The bone fixation plate of any of claims 1-10, wherein the overall stiffness of the bone fixation plate is selected based on a virtual surgical model including an underlying bone and a surrounding tissue with their corresponding mechanical properties.

12. The bone fixation plate of any of claims 1-11, wherein the bone fixation plate includes an outer solid shell extending between the first surface and the second surface along a periphery of the bone fixation plate.

13. The bone fixation plate of any of claims 1-12, wherein the plurality of screw holes are threaded and countersunk with respect to the first surface.

14. The bone fixation plate of any of claims 1-13, wherein an internal porous core defined by the network of interstices includes a pore geometry to modulate stiffness, wherein the pore geometry includes orthogonally oriented struts of 300-micrometer diameter interspersed by 300- micrometer pores.

15. The bone fixation plate of any of claims 1-14, wherein the bone fixation plate comprises a shape memory alloy, such as a NiTi alloy (e.g., nitinol).Attorney Docket No. 103361-579WO1T2024-04616. A system comprising: a bone fixation plate comprising: a first end and a second end opposite and spaced apart from the first end along a longitudinal axis; a first surface extending from the first end to the second end; a second surface opposite from the first surface; a network of interstices disposed between the first surface and the second surface; and a plurality of screw holes defined by and extending through the bone fixation plate from the first surface to the second surface; at least one manufacturing device configured to produce at least one portion of the bone fixation plate; and a control system in communication with the manufacturing device, the control system comprising computer-readable instructions for producing the bone fixation plate, the computer- readable instructions being based on an anatomical model of a patient.

17. The system of claim 16, wherein the at least one manufacturing device comprises an additive manufacturing device configured to produce the bone fixation plate and the network of interstices therein according to the computer-readable instructions of the control system.

18. The system of claim 17, wherein the additive manufacturing device produces the bone fixation plate having a desired stiffness according to the computer-readable instructions of the control system.

19. The system of any of claims 17-18, wherein the additive manufacturing device produces the bone fixation plate further having a desired shape or a desired curvature according to the computer-readable instructions.

20. The system of claim 17, wherein the at least one manufacturing device further comprises a bending manufacturing device configured to adjust a shape and / or a curvature of the bone fixation plate according to the computer-readable instructions of the control system.Attorney Docket No. 103361-579WO1T2024-04621. The system of claim 16, wherein the at least one manufacturing device comprises a bending manufacturing device configured to adjust a shape and / or a curvature of the bone fixation plate according to the computer-readable instructions of the control system.

22. The system of any of claims 16-21, wherein the network of interstices is fabricated, via the manufacturing device, to have a pattern matching the desired stiffness of the bone fixation plate.

23. The system of any of claims 16-22, wherein the manufacturing device is a metal-powderbased 3D printing device.

24. The system of any of claims 16-23, wherein the network of interstices comprises a network of pores, each of the pores in the network of pores having either a uniform shape and size or a non-uniform shape and size.

25. The system of any of claims 16-24, wherein the bone fixation plate further comprises a plurality of struts extending between at least a portion of the bone fixation plate between the first surface and the second surface, the plurality of struts defining the network of interstices.

26. A method of fabricating a bone fixation plate, the method comprising: providing an anatomical model of a patient; calculating a stiffness of a portion of the anatomical model and the curvature of a portion of the anatomical model; calculating a set of computer-readable instructions for producing the bone fixation plate having a shape matching the stiffness and the curvature of the portion of the anatomical model; fabricating, via a manufacturing device configured to receive the set of computer- readable instructions, the bone fixation plate with a network of interstices providing a stiffness of the bone fixation plate that substantially matches the stiffness of the portion of the anatomical model.

27. The method of claim 26, wherein the manufacturing device is an additive manufacturing device comprising a metal-powder material.Attorney Docket No. 103361-579WO1T2024-04628. The method of any of claims 26-27, wherein a plurality of struts define the network of interstices of the bone fixation plate, the plurality of struts being arranged orthogonally to each other.

29. The method of any of claims 26-28, further comprising: implanting the bone fixation plate on an anatomical feature of the patient on which the anatomical model was based by installing a screw through a screw hole extending through the bone fixation plate.

30. The method of any of claims 26-29, wherein the bone fixation plate is configured to provide sufficient support to a portion of a patient’s anatomy and to provide sufficient flexibility to reduce stress fractures of the bone fixation plate.

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