Melted electrowriting tissue scaffold

Abstract

A fused electrowriting soft tissue stent comprising: a first region having one or more groups of fibers and a second region having one or more groups of fibers; and an interface region joining the first region and the second region, the interface region being electrically written with the first region and the second region in a continuous print path such that fibers in the interface region are each joined between respective pairs of fibers in the first region and the second region.

Description

Technical Field

[0001] The present disclosure generally relates to implantable devices or stents, such as those used for engineered heart valves. More particularly, the present invention relates to interfaces between structural regions and gradients within a stent.

[0002] Incorporation of References

[0003] The inventors previously developed a new method of providing a stent or implant using molten electrowriting (MEW), as disclosed in Patent Cooperation Treaty application PCT / AU2020 / 210877, the contents of which are incorporated herein by reference. Background Art

[0004] Biological tissues are complex, multiphasic, heterogeneous, and layered structures that exhibit exquisite properties that are well-suited to their functions. Furthermore, tissue interfaces not only play a critical role in physically connecting heterogeneous regions but also exhibit gradients in structural, cellular, and mechanical characteristics, crucially influencing the overall function of a tissue or organ. Therefore, interfaces represent a critical yet challenging problem when attempting to biofabricate scaffolds for these tissues.

[0005] Most research on tissue interface engineering involves soft-hard tissue interfaces in fields such as ligaments and tendons, cartilage, and dental and craniomaxillofacial implants. In these applications, heterogeneous scaffolds are realized through advanced manufacturing techniques such as multiaxial extrusion, varying degrees of cross-linking, two-step phase separation, multimaterial bioinks, and controlled spatial deposition of biomaterials through advanced three-dimensional (3D) printing technologies.

[0006] Fibrous scaffolds, especially electrospun mesh scaffolds, have been extensively studied in the field of soft tissue engineering (e.g., skin, nerve, vascular, or cardiac tissue). However, comparable studies on interface and gradient design are still insufficient. Interface construction between regions is mainly achieved in a layer-by-layer manner, such as by changing printing parameters or solution concentration during the manufacturing process, subsequent cross-linking, or by individual layer-by-layer assembly. The lack of interface complexity and gradient structure in electrospun scaffolds may be due to the lack of precise control of fiber orientation when manufactured using this technology.

[0007] Melt electrowriting (MEW) is a high-precision additive manufacturing technology capable of printing complex fiber scaffold structures with submicron resolution. The complexity and resolution characteristics of MEW have enabled the production of functional biomimetic soft tissue scaffolds for skin, nerves, myocardium, cartilage, and aortic heart valves. Similar to electrospinning, MEW has achieved interface construction on a layer-by-layer basis. Recently, studies have demonstrated the ability to use MEW technology to prepare heterogeneous fiber structures within a single layer, marking a major step forward in the field of fiber scaffold manufacturing. Despite this, the concept of heterogeneous MEW printing is still in its infancy, and therefore the design of interfaces between heterogeneous regions has not been a major focus of previous research.

[0008] It is acknowledged that, if any prior art is referred to herein, such reference does not constitute an admission that the prior art forms part of the common general knowledge in the art in Australia or any other country. Summary of the Invention

[0009] In a first aspect, a melt electrowriting support is disclosed. The support comprises a first region having one or more groups of fibers and a second region having one or more groups of fibers. The first and second regions are connected by an interface region, which is electrowritten along a continuous print path within the same layer as the first and second regions, such that fibers in the interface region are connected between corresponding fiber pairs in the first and second regions.

[0010] Scaffolds can be used in the engineering of biodegradable or non-biodegradable implantable devices.

[0011] The first region and the second region may be heterogeneous in that they differ from each other in one or more spatial parameters.

[0012] The path of the at least one fiber within the interface region may be defined manually or by a mathematical function.

[0013] The function may be such that the fiber has a complex shape as it transitions from one of the first region or the second region through the interface region to the other of the first region or the second region.

[0014] The second region may have a higher porosity than the first region.

[0015] The higher porosity in the second region can be formed at least in part by connecting every two or more adjacent fibers in the first region as they transition into the second region.

[0016] The higher porosity in the second region can be formed at least in part by causing the fibers to fan outward as they transition into the second region.

[0017] The higher porosity in the second region can be formed at least in part by arranging the continuous printing path so that another set of fibers in the first region are deposited offset from the other fibers by a distance less than the pore size.

[0018] The first region or the second region or both may include a first set of fibers arranged generally parallel to one another and a second set of fibers arranged generally parallel to one another, the second set of fibers being arranged at an angle relative to the first set of fibers and preferably arranged transversely, each fiber in the second set of fibers having a serpentine arrangement having defined valleys and peaks.

[0019] The fiber arrangement in the interface region can be at least partially biomimetic. For example, the fiber arrangement can be based on a simplified form of collagen fiber organization observed in the interface region of a biological sample of soft tissue. The collagen fiber organization can be determined by orientation analysis.

[0020] The scaffold may have multiple layers of fibers.

[0021] A portion of the stent may have a different number of layers than another portion of the stent. The stent may be a heart valve stent. The first region may be a leaflet, and the second region may be an interleaflet triangle.

[0022] On the other hand, disclosed herein is a method for providing a melt electrowriting soft tissue scaffold having at least two structurally heterogeneous regions and an interface region therebetween, the method comprising providing a printing path along which a polymer melt material is continuously extruded during a melt electrowriting process to form the heterogeneous regions and the interface region.

[0023] The method may comprise causing the continuous printing path to be defined in segments such that it is defined by different functions in the heterogeneous region and the interface region.

[0024] In another aspect, a method for designing a MEW scaffold for soft tissue having multiple structural regions is disclosed. The method includes: imaging the soft tissue using an imaging modality capable of imaging collagen fiber structure within the soft tissue; identifying an interface region from the structural regions; applying an orientation analysis to at least an image from the imaging step that shows the interface region; determining a simplified form of the fiber distribution observed in the orientation analysis, and configuring a printing path for an interface gradient region of the scaffold corresponding to the interface region of the soft tissue based on the simplified form of the fiber distribution.

[0025] The one or more functions defining the continuous printing path are mathematical functions, preferably providing a curved or serpentine shape as the continuous printing path traverses the interface gradient region.

[0026] The present invention has scaffold-based biomedical applications because natural tissues are rarely homogenous, ie gradients exist in all tissue types.

[0027] In one aspect of the invention, a scaffold comprises a gradient of fibers comprising the scaffold. A gradient refers to a gradual change in a property over space. This change can be in terms of geometric characteristics (e.g., size, thickness, shape, density, orientation), mechanical properties (rigidity, elasticity), composition (cell / tissue type), chemical properties (pH), or more. The scaffolds of the present invention can be used to generate multiple gradient properties in a single scaffold.

[0028] For example, disclosed herein are melt-electrowritten stents comprising gradient porosity. The terms "gradient porosity" or "porosity gradient" refer to a change in the size and / or shape of pores within the stent over the entire area of ​​the stent. The pore size can gradually increase or decrease spatially along one or more dimensions of the stent. The porosity gradient is determined by a mathematical function, where the shape of the pores within the stent can be, for example, rectangular, diamond-shaped, or serpentine.

[0029] Preferably, the scaffold comprises multiple layers of fibers. The multiple layers of fibers may comprise gradients of geometrical characteristics and / or mechanical properties to form a gradient scaffold. Each fiber of the scaffold may comprise a gradient of geometrical characteristics and / or mechanical properties to form a gradient scaffold. For example, the fibers may be thicker, denser and / or harder in one region and gradually become thinner, less dense and / or more flexible in another region. The one or more functions defining the continuous printing path across the gradient region are mathematical functions that define the geometrical characteristics and / or mechanical properties of the fibers in that region.

[0030] The stent gradient design can be applied to all or part of the stent area.

[0031] Gradient functions can be generated for heart valve applications, incorporating the patient-specific anatomy of the heart valve inspired by the orientation, distribution, and density of native collagen fibers.

[0032] The scaffold fibers forming the porosity gradient can be intentionally aligned parallel or orthogonal to the interface between the two regions. In the case of a heart valve, this would mean that they are aligned with the interface between the leaflets and the interleaflet triangles / commissures / annulus.

[0033] The porosity in the gradient can be arranged to be smaller in the region of the commissures and larger in the belly of the leaflet, or vice versa, to reflect areas of higher load.

[0034] The fiber density gradient can be arranged to be higher in the area of ​​the commissures and belly of the leaflet above the annulus and lower inside the leaflet, or vice versa, to reflect areas of higher loading.

[0035] Any of the above types of gradients (geometrical features, mechanical properties, composition, chemical properties, etc.) can be used in combination or individually to design fused electrowritten stents for heart valves or other tissues. BRIEF DESCRIPTION OF THE DRAWINGS

[0036] Embodiments will now be described, by way of example only, with reference to the accompanying drawings.

[0037] Figure 1 Schematic diagram of the components of the melt electrowriting (MEW) equipment.

[0038] Figure 2 An example of a digital microscope image of a MEW scaffold with two heterogeneous regions and a continuous interface between the regions.

[0039] Figure 3-1 Schematic diagram of a scaffold with uniform porosity distribution.

[0040] Figure 3-2 for Figure 3-1 Schematic diagram of the arrangement of fibers in a scaffold connected in pairs to expand pore size.

[0041] Figure 3-3 Schematic diagram of "halving" a scaffold: the pore size of a certain part of the scaffold is halved by adding another fiber layer to the scaffold.

[0042] Figure 3-4 Schematic diagram of the "fan-shaped expansion" of the scaffold: the fibers in the second (right) area are distributed in a fan-shaped manner to gradually increase the pore size in this area.

[0043] Figure 3-5 An example image of the MEW manufacturing apparatus showing a section of a stent being "halved."

[0044] Figure 3-6 An example of a scaffold where fibers fan out in one section to reduce pore size.

[0045] Figure 3-7 An example of a scaffold where fibers fan out in one section to increase pore size.

[0046] Figure 4-1 Schematic diagram of overlapping interface, stitched interface and continuous interface.

[0047] Figure 4-2 Digital microscopy (top) and scanning electron microscopy (SEM) (bottom) images of MEW polycaprolactone (PCL) scaffolds fabricated using overlapping, stitching, and continuous interface methods. The pore patterns are 1 mm square and 0.5 mm diamond, respectively. Scale bar = 1 mm.

[0048] Figures 5-1 to 5-3 Images of uniaxial tensile tests of biphasic MEW scaffolds. The interface methods are overlapping, sutured, and continuous techniques. Figure 5-1Time-lapse images showing the tensile testing process of a continuous interface 1 mm pore serpentine and 0.5 mm pore square scaffold. Scale bar = 5 mm. Figure 5-2 A typical stress-strain curve of a continuous interface scaffold is shown, with regions marked as i) elastic deformation, ii) plastic deformation, and iii) scaffold fracture. Figure 5-3 The comprehensive stress-strain curves of different interface methods of serpentine-square scaffolds and the control group are shown.

[0049] Figure 5-4 The Young's modulus, yield strength, and ultimate tensile strength of biphasic scaffolds with different interface methods and a control scaffold are shown (n=3, mean values, error bars are standard deviations, and significant differences were assessed using one-way analysis of variance (ANOVA) with Tukey's multiple comparison test. Statistical significance is indicated on a single bar when compared to every other bar on the graph, and the data are at least significantly different, *p≤0.05, **p≤0.01, ***p≤0.001, ****p≤0.0001).

[0050] Figure 6-1 Images showing clamping distances of 2.5 mm, 4.5 mm, and 6.5 mm applied to MEW scaffolds to test uniaxial tensile properties.

[0051] Figure 6-2 Representative stress versus strain plots are shown for strain applied to a continuously bonded serpentine square scaffold structure at clamping distances of 2.5 mm, 4.5 mm, and 6.5 mm, respectively.

[0052] Figure 6-3 Shown are the Young's moduli measured with applied strain at different clamping distances.

[0053] Figure 6-4 Shown are the yield strengths measured under applied strain at different clamping distances.

[0054] Figure 6-5 Shown are the ultimate tensile strengths measured with applied strain at different clamping distances.

[0055] Figure 7-1 The flexure test apparatus is schematically depicted, where θ represents the measured bending angle.

[0056] Figure 7-2 Close-up view of the biphasic scaffold showing the location of the pivot axis, scale bar = 1 mm.

[0057] Figure 7-3 is a test image of a multiphase scaffold with a serpentine pattern, scale bar = 10 mm.

[0058] Figure 7-4Flexural stiffness data are depicted (mean of n = 3, error bars indicate standard deviation, two-way ANOVA with Tukey's multiple comparison test resulted in p < 0.0001 for pattern effect, p = 0.0002 for interface effect, and non-significant differences between individual means).

[0059] Figure 8-1 is a sample graphical user interface for generating continuous and spatially heterogeneous G-code.

[0060] Figure 8-2 is an example scaffold designed using the graphical user interface (GUI).

[0061] Figure 8-3 is another example bracket designed using a graphical user interface (GUI).

[0062] Figure 8-4 Digital microscope images are shown, including Figure 8-2 Images of the designed MEW PCL scaffold shown, high-resolution image of the auxetic star pattern with 1 mm pores in part “A” and high-resolution of the scalloped interface showing a continuous transition from 1 mm pores to 2 mm pores.

[0063] Figure 8-5 Digital microscope images are shown, including Figure 8-3 Images of the designed MEW PCL scaffold structure, with a high-resolution image of the diagonal serpentine pattern of 1 mm pores in section “A,” and a high-resolution image of the scalloped interface showing a continuous transition from 1 mm pores to 2 mm pores. Scale bar = 2 mm.

[0064] Figure 9-1 This is an image showing the aortic valve region.

[0065] Figure 9-2 Depicted are the 3D reconstructed valve structures and cutting planes viewed from the i) coronal (external) direction, ii) sagittal direction, iii) axial direction, and iv) coronal (internal) direction.

[0066] Figure 9-3 is a second harmonic generation (SHG) image of the aortic valve showing the commissures and interfaces with the beginnings of the adjacent leaflets.

[0067] Figure 9-4 is an SHG image of a slice of an aortic valve sample, showing the commissures.

[0068] Figure 9-5 is an SHG image of a sample slice of an aortic valve showing the interleaflet triangles.

[0069] Figure 10 Schematic diagram of the hybrid immersion fixation device used to fix porcine aortic valve tissue samples.

[0070] Figure 11-1 is a second harmonic generation (SHG) image of the aortic valve commissure, showing the location of the imaging plane selected for correlative focused ion beam scanning electron microscopy (FIBSEM); scale bar = 1 mm.

[0071] Figure 11-2 is a low-magnification FIBSEM image taken in the circumferential direction showing the aortic valve fiber cell microstructure in the upper region of the commissures, scale bar = 10 μm.

[0072] Figure 11-3 is Figure 11-2 Magnified images of the area bounded by the dashed box in , showing a cross section of a single collagen fiber running in the circumferential (C) direction, and cross sections of longitudinal fibers running in the radial (R) and longitudinal (L) directions, scale bar = 1 μm.

[0073] Figure 12-1 Color-mapped orientation analysis of SEG images of interlobular triangular regions is depicted, scale bar = 0.5 mm.

[0074] Figure 12-2 is a simplified schematic diagram of the orientation of collagen fibers in the interlobular triangles.

[0075] Figure 12-3 Color-mapped orientation analysis of the resulting MEW scaffold and interfacial region is depicted, scale bar = 2 mm.

[0076] Figure 12-4 Overlapping orientation distributions from interlobular triangular tissue samples and MEW scaffolds are shown.

[0077] Figure 12-5 The resulting MEW bracket is shown labeled with the input G-code showing the design parameters.

[0078] Figure 12-6 Snapshots of the G-code path at times = 1, 2, 6, 29, and 74 seconds after the start time with corresponding time stamps are shown to show the continuous printing path at these times.

[0079] Figure 12-7 is an SEM image showing a close-up view of the continuous fiber interface using the joining technique, where the fibers have fused along the joining path, scale bar = 100 μm.

[0080] Figure 13-1 is an image of the biaxial tensile testing apparatus used to apply strain in the circumferential and longitudinal test directions, scale bar = 5 mm.

[0081] Figure 13-2Graphs showing the measured Young's modulus, yield strength, and yield strain for circumferential and longitudinal tests.

[0082] Figure 13-3 The hysteresis behavior from cyclic testing at constant strain is depicted, showing representative stress-strain diagrams in two directions.

[0083] Figure 13-4 Quantification of hysteresis, shown as the area under the unloading curve divided by the area under the loading curve for the same cycle, is depicted.

[0084] Figure 13-5 Representative relaxation behaviors after four incremental increases in strain (2.5% and 5% in the longitudinal and circumferential directions, respectively) held for 1000 s are depicted, and the percent relaxation at each step was quantified (all data show the mean of n = 3, error bars represent one standard deviation, significant differences were assessed using a parametric ratio paired t-test, *p ≤ 0.05, **p ≤ 0.01, ***p ≤ 0.001).

[0085] Figure 13-6 The relaxation percentage is plotted as a function of the number of applied strain steps.

[0086] Figure 14-1 is a conceptual illustration of an example of a sinusoidal interface.

[0087] Figure 14-2 is a conceptual illustration of an example of an angled interface.

[0088] Figure 14-3 is a conceptual illustration of an example of an arrow-shaped interface.

[0089] Figure 14-4 An example scaffold schematic for a layer is shown, illustrating a sinusoidal interface transitioning between regions with a rectangular grid pattern and regions where fibers are arranged in an "auxetic star" pattern.

[0090] Figure 14-5 An example scaffold schematic for a layer is shown, illustrating an "angled" interface transitioning between regions with a rectangular grid pattern and regions where fibers are arranged in an "auxetic star" pattern.

[0091] Figure 14-6 An example scaffold schematic for a layer is shown, showing arrows or arrow-like interfaces transitioning between regions having a rectangular grid pattern and regions where fibers are arranged in an "auxetic star" pattern.

[0092] Figure 15-1 Shows the Figure 14-4 Parameters of the scaffold layer schematic shown in .

[0093] Figure 15-2 Shows the Figure 14-5 Parameters of the scaffold layer schematic shown in .

[0094] Figure 15-3 Shows the Figure 14-6 Parameters of the scaffold layer schematic shown in .

[0095] Figure 16 Schematic diagram showing exemplary gradient porosity scaffolds and the underlying mathematical function controlling the gradient porosity, where the values ​​of the initial pore size (p0) and the gradient coefficient (m) can be varied. The gradient function has been applied to three different patterns (rectangular, diamond, or serpentine) and affects the pore size parallel or orthogonal to the gradient direction (it is applied horizontally from left to right).

[0096] Figure 17 This figure shows an example preview of parallel and orthogonal gradient porosity scaffolds with rectangular, diamond, or serpentine patterns. All scaffold structures have dimensions of 25 × 25 mm, an initial pore size (p0) of 0.5 mm, and gradient coefficients (m) of 0.5, 1.0, or 1.5. Note that in the case of orthogonal gradient porosity scaffolds, the gradient coefficient refers to the steepest gradient line, and the remaining lines are interpolated between that gradient line and flat (zero gradient).

[0097] Figure 18 Shown are exemplary gradient porosity scaffolds fabricated using melt electrowriting using a 23G needle, 100 kPa pressure, 3 mm working distance, 3.8 kV potential difference, 400 mm / min printing speed, 75°C syringe temperature, 85°C needle temperature, and 30°C bed temperature. Each scaffold was made of polycaprolactone, measured 25 × 25 mm, had five layers, and had an average fiber diameter of 20 μm.

[0098] Figure 19 Shown are local strain maps of a gradient scaffold undergoing constrained biaxial tension testing. The scaffold was clamped using a 15 mm clamp and strained at 1% / s in the horizontal direction to 100% strain while being fixed in the vertical direction. Images were taken at 20% strain. The color map shows the uneven distribution of the principal engineering strains calculated using VIC-2D software (Correlated Solutions, USA).

[0099] Figure 20 Heterogeneous loading in different regions of the valve is shown, particularly in the mid-ventral and commissural regions. Figure adapted from: Emmert et al., Science Translational Medicine 10, No. 440 (2018). https: / / doi.org / 10.1126 / scitranslmed.aan4587.

[0100] Figure 21 Aspects of exemplary heart valve stent designs incorporating gradients are shown. A) Designs from PCT / AU2020 / 210877 have uniform serpentine amplitude / wavelength throughout, aligned in only two directions: circumferential (blue, left-right) or radial (red, up-down) and anisotropic fiber density / porosity (higher density circumferentially, lower density radially). B) Radial fibers cross orthogonal (or nearly orthogonal) to the leaflet lines. C) Circumferential fibers move somewhat parallel to the leaflet lines. D) Serpentine wavelength and / or amplitude varies across the leaflet. E) Fiber density varies throughout the leaflet (in higher loaded areas, such as the commissures and base of the leaflet belly).

[0101] Figure 22 is a schematic diagram of an exemplary heart valve stent design incorporating gradient porosity.

[0102] Figure 23 is a schematic diagram of a method for continuously and gradually combining fibers between the three leaflets. Note that for simplicity, line 0 is drawn as a straight line, but this may involve serpentine, gradient, and / or other patterns. DETAILED DESCRIPTION

[0103] In the following detailed description, reference is made to the accompanying drawings, which form a part of the detailed description. The illustrative embodiments described in the detailed description depicted in the accompanying drawings are not intended to be limiting. Other embodiments may be utilized, and other changes may be made without departing from the spirit or scope of the subject matter presented. It will be readily understood that the various aspects of the present disclosure, as generally described herein and illustrated in the accompanying drawings, may be arranged, substituted, combined, separated, and designed in a variety of different configurations, all of which are contemplated in the present disclosure.

[0104] Gradient structures are abundant in living tissues and play a key role in their function. The ability to fabricate structures that replicate these gradients and interfaces could be beneficial in achieving functional constructs for tissue engineering applications.

[0105] Gradient scaffolds have been studied in the context of tissues such as bone, tendon, vasculature, and myocardium, to name a few. These scaffolds have been fabricated using techniques that spatially control the properties of the material (such as temperature, pH, or concentration) throughout the material. Fiber fabrication techniques such as MEW have also been applied to the creation of gradients, however, only for simple rectangular geometries. Furthermore, gradient scaffolds can be used to create disease or tissue models, for example using gradient stiffness hydrogels to elicit graded cellular responses.

[0106] New methods for fabricating microfibrous scaffolds, such as MEWs, can enable more complex interface or gradient designs. MEWs have the ability to achieve complex and customizable interfaces and gradient structures, which may benefit a wide range of tissue engineering applications or disease modeling.

[0107] Melt electrowriting (MEW) is a highly precise additive manufacturing technology that has the potential to leverage specific aspects of microstructural features observed in natural tissues, such as collagen orientation, into the design of complex biomimetic scaffolds. Printing of MEW scaffolds relies on precise control of five main process parameters: temperature, pressure, voltage, working distance, and collector speed, resulting in a highly controllable jet of molten polymer with micron resolution that is deposited on a collector in a direct-write mode. However, unlike other 3D printers, MEW does not allow for stopping and starting extrusion during a print run, as this would disrupt the Taylor cone and lead to printing defects. Therefore, conventionally, MEW printing is performed for each region of similar spatial features in one uninterrupted print run. Tissues with spatially heterogeneous regions are therefore fabricated by printing each region separately and then joining them together, for example by suturing.

[0108] This paper discloses an inventive method for designing continuous, user-defined interfaces between separate structural regions of implantable devices using MEWs, where the interface regions and structural regions are printed continuously on each layer. The design strategy for joining the different architectures of a multiphase MEW scaffold requires careful consideration, as this may affect its ultimate mechanical and biological functionality.

[0109] The continuous interface disclosed herein is applied to MEW printing of various implantable devices. An exemplary application is the fabrication of MEW stents for aortic valves.

[0110] The complexity and heterogeneity of the aortic valve make it a good demonstration for examining the technical capabilities of MEW in the fabrication of complex scaffolds for soft tissue interfaces. However, it should be understood that this approach is applicable to engineering other tissues besides heart valve scaffolds.

[0111] The average human aortic valve experiences over 30 million cycles per year, totaling over 2 billion cycles over a 70-year lifespan. This remarkable hemodynamic property is achieved through the combined biomechanical behavior of each structurally distinguishable region of the valve and the interfaces between them working together. The inventors previously demonstrated a MEW scaffold with mechanical properties similar to those of physiological heart valve leaflets using a bioinspired design approach. This was achieved using available data on the relationship between the mechanical properties of the surrounding leaflets and the collagen recruitment mechanisms during loading.

[0112] In one application, the disclosed method is relevant to regions of the heart valve beyond the leaflets and the interfaces between them, potentially unlocking further applications of MEW to enhance valve function. Regions such as the commissures, interleaflet triangles, and the interfaces between them all play a key role in the biomechanical behavior of heart valves.

[0113] Heart valves are heterogeneous structures with highly diverse structural and mechanical properties. These gradient properties exist not only within the heart valve leaflets, but also in surrounding structures such as the commissures, interleaflet triangles, and annulus, each of which plays an important role in valve function. Therefore, in the context of engineered heart valves, the ability to generate gradient scaffolds and thereby control local mechanical properties is extremely valuable.

[0114] Therefore, the present invention applies the concept of gradient stent to the design of fiber heart valve stent ( Figure 22 、 23 ).

[0115] PCT / AU2020 / 210877 identifies the benefits of continuously connecting fused electrowritten fibers between leaflets to improve printing and flexure performance. However, in previous disclosures, the orientation changes between regions, while continuous, were abrupt. The present invention provides a way to gradually and continuously combine fibers within and between regions.

[0116] Figure 1 Components of a MEW apparatus 10 are schematically shown, including a nozzle 12 for extruding a polymer melt onto a collector 14. For clarity, the "x", "y" and "z" directions are used herein with reference to Figure 1 The orientation shown in FIG. The marking of the directions may be different and does not limit the scope of the invention.

[0117] The nozzle 12 is positioned above the collector 14 at a working distance (height), defined here in the "z" direction. As the nozzle 12 extrudes the polymer, relative movement between the nozzle 12 and the collector 14 in the "x" and "y" directions, or a combination thereof, enables the formation of a 2D MEW scaffold in layers; multilayer 3D structures can be created by increasing the relative motion between the nozzle and the collector along the "z" direction. Translation of the nozzle 12 relative to the collector 14 in the "y" direction (or vice versa) defines the translation speed.

[0118] The polymer jet extruded from the nozzle 12 is deposited as fibers 16 onto a collector 14. The fibers 16 follow a predefined printing path to form the desired structure. At the end of the MEW printing process, the scaffold will have multiple layers of fibers. The number of layers will depend on the MEW setup and the desired configuration of the scaffold.

[0119] Some, but not necessarily all, scaffolds that can be manufactured using the methods described herein are heterogeneous. A heterogeneous scaffold comprises at least two regions of spatially heterogeneous structure. Spatial heterogeneity can be a difference in one or more parameters, such as, but not limited to, pore size, one or more of size, fiber orientation and fiber density, pattern, length, etc.

[0120] According to the present invention, adjacent structural regions are each formed continuously at one or more layers of the stent with an interface therebetween.

[0121] Figure 2 is a digital microscope image of a MEW scaffold layer 20. Layer 20 includes two structural regions 22, 24 and a continuous interface region (which may also be simply referred to as an "interface") 26 joining the regions 22, 24. In this example, the two regions 22, 24 are spatially heterogeneous. However, the interfaces described herein can also be used to join regions that are not spatially heterogeneous.

[0122] In this example, region 22 has a larger pore size than region 24, and the pores in region 22 are more rectangular in shape, while the pores in region 24 are more rhombus-like in shape. A continuous interface 26 exists between the heterogeneous regions 22 and 24. The exact configuration of the fibers shown does not represent an essential element of the present invention. Rather, they are provided merely as examples to illustrate the concept of a "continuous interface" as intended in this disclosure, i.e., an interface joining two structural regions that is printed in a continuous manner on the same layer along with the structural regions.

[0123] For example, from Figure 2 As can be seen in Figure 2, starting from region 22, the fibers follow the path of the defined grid pattern in region 22, but then deviate at the boundary between region 22 and the interface region 26 to follow the interface path controlled by the interface function as the fibers enter the interface region 26. At the boundary between the interface region 26 and region 24, the fibers deviate again from their current path (the interface path) to follow the path of the diamond pattern in the defined region 24. This deviation between the path of the defined region pattern and the interface path occurs throughout the print to print the entire support layer in one continuous print. Therefore, the entire print path remains continuous, but at the region boundaries with the interface region, it switches to follow different paths defined for each region. This arrangement contrasts with conventional MEW printing, in which different structural regions are printed separately to avoid deviations in the print path prepared for the support. These regions are then joined together, typically by stitching or by overlapping or by a combination of both.

[0124] In this method, during printing, for each layer, the fibers of that layer are formed by depositing polymer along a printing path that traverses the layer 20 multiple times until the layer is formed, and then the MEW printer continues to form additional layers. The printing of additional layers can follow the same printing path as the first layer, or different printing paths can be followed if the scaffold to be manufactured is configured to have variations across the layers (i.e., across the thickness of the sheet along the "z" dimension).

[0125] A print path can also be considered to include multiple round trips. The number of passes can include any one or a combination of the following: one or more times across all or part of the width of a layer (e.g., along the "y" direction); one or more times across all or part of the height of a layer (e.g., along the "x" direction); one or more times across all or part of the tilt direction of a layer.

[0126] More generally, the print path will be determined by the MEW printing algorithm (e.g., written in G-code) that defines the device being manufactured (e.g., a stent). A "continuous print path" as used herein means the path along which the polymer is continuously extruded. The times used to print an entire sheet of paper will be part of a continuous print path. The print path can include curved and straight segments, again depending on what is being manufactured.

[0127] Therefore, a continuous print path will pass through different regions. To facilitate this, the algorithm defining the paths used for different regions will switch between these paths when the fiber encounters a boundary between two regions, so as to form a continuous print path that transitions between the two regions in a layer without pausing the MEW printing process for that layer.

[0128] In some embodiments, the print path indicating the fiber deposition location or path in the interface region 26 will be an interface function, meaning a single or possibly composite mathematical function that defines a particular shape while still allowing the boundaries of the regions to transition continuously into one another, rather than simply connecting between fiber locations in regions 22, 24. Examples include, but are not limited to, mathematical functions that form sinusoidal, parabolic, arrow, and angled shapes. Conceptual illustrations of examples of sinusoidal, angled, and arrow-shaped interfaces 1401, 1402, and 1403 are provided in FIG. Figure 14-1 、 Figure 14-2 and Figure 14-3 Shown in. Figure 14-4 、 14-5 14-6 show example scaffold schematics of layers with sinusoidal, "angled," and "arrowhead" shaped interfaces, respectively, transitioning between regions with a rectangular grid pattern and regions where fibers are arranged in an "auxetic star" pattern. Figure 14-4 、 14-5and 14-6 bracket schematics are shown in the Figure 15-1 、 15-2 and 15-3.

[0129] The shape of the interface can be manually defined instead of being defined by a function or composite function (i.e., "function-defined"), or in addition to being defined by a function or composite function. For example, there can be a combination of function-defined and manually defined shapes for fibers within one or more layers. The interface fibers in one or more layers can all be manually defined. The interface fibers in one or more layers can all be function-defined.

[0130] Furthermore, depending on the fiber arrangement to be provided for the regions connected by the interface, the different print paths prepared for the fibers of different interfaces can have different shapes or interface functions, each print path extending between a corresponding pair of fibers from adjacent structural regions. In this case, it is preferred that the different interface functions defining a series of adjacent print paths in the interface region define a shape that changes gradually rather than abruptly.

[0131] The choice of interface shape or function will depend on various factors, such as the desired printing time and any mechanical properties desired for the interface region or the overall tissue scaffold structure (e.g., specific bending or flexing requirements). The choice may also depend on the values ​​of the MEW process parameters. For example, if any fusion or bonding is intended to occur between an already deposited fiber portion and a fiber portion deposited later (e.g., during a return pass), a shorter time may be required to traverse the interface region (thus requiring a shorter print path).

[0132] A continuous interface for joining two structural regions (further wherein the interface is defined by functionality as described above) has been demonstrated in experiments conducted by the inventors to provide a maximum degree of flexure compared to conventional techniques for joining two structures. This is useful, particularly for applications such as heart valve stents, where flexure or bending in an implanted valve is required to function in a hemodynamic environment, providing the opportunity to manufacture a valve that can better maintain its performance over time.

[0133] Alternatively, two or more printing paths can be defined so that the fibers deposited along the paths are immediately adjacent to each other, forming bundles. Under the correct MEW settings, the bundles can fuse or bond to each other. This allows for the fabrication of scaffold sections that mimic anatomical features found in biology. For example, the inventors observed that in test porcine heart valve samples, collagen fiber bundles split and also connected individual collagen fibers into larger bundles. This is described later in this specification.

[0134] By providing an interface through which two structural regions transition between each other, pore size can be varied between different parts of a printed structure. Three techniques for varying the pore size between these parts are proposed: connecting, halving, and fanning. The same techniques can also be applied to varying the porosity within individual structural regions, either instead of or in addition to transition regions (i.e., interfaces) between different regions.

[0135] Figures 3-1 to 3-4 Conceptual depiction of the concepts of "connection", "halving" and "fanning out". Figure 3-1 Schematically shows a stent with a consistent pore size that does not vary between the left-hand and right-hand portions of the stent. Figure 3-2 In the example of "joining", each pair of adjacent fibers from the left-hand part is joined together, resulting in a doubling of the pore size on the right-hand side. This can be a way to create a transition from one structural area to a second structural area with a larger pore size. This will also result in doubling the number of fiber layers in the second area. This technique reflects observations made from natural tissue (see, for example, Figure 9-5 ), where small collagen bundles gradually combine to form thicker bundles, as observed within interfaces elsewhere in the body. The concept of "coupling" can be summarized as the joining of two or more fibers. Potentially, different numbers of fibers can be joined at different areas in the scaffold during this process to produce joined bundles of different sizes.

[0136] The halving method deposits additional scaffold layers in the first (left-hand) section at an offset equal to half the pore size. This results in the total pore size in the first section being halved while maintaining equal fiber layers between sections. Figure 3-3 The halving of the pore size with the addition of an additional fiber layer 32 is schematically depicted. Figure 3-5 Example images of this happening in a MEW-fabricated device are provided.

[0137] The "halving" method can be generalized to reduce the aperture in the region where the additional layer is deposited by an amount other than half. For example, "halving" can be performed twice, with a 1 / 3 aperture offset and a 2 / 3 aperture offset, to reduce the aperture to 1 / 3 of the original aperture. Therefore, the "halving" concept can be generalized to the "division" or "fraction" concept.

[0138] The fanning method causes a gradient change in porosity. The gradient change can be set to occur over a "fanning height" that is predefined or selected by the designer of the MEW scaffold. Fanning can occur to gradually reduce the pore size ( Figure 3-4 、 Figure 3-6 ) or gradually increase the pore size ( Figure 3-7 ).

[0139] experiment

[0140] Example 1

[0141] Comparison of continuous bonding and other methods of bonding multiphase MEW scaffolds

[0142] The effects of the design of the interfacial region on the behavior of the printed scaffolds were systematically investigated. In addition to the above-mentioned continuous interfacial bonding, two other interfacial bonding methods were also investigated, including overlapping and stitching. Figure 4-1 , which includes schematic diagrams of continuous interfaces, stitched interfaces, and overlapping interfaces.

[0143] Three different patterns (square, diamond, and serpentine) were selected to combine using three interfacing printing strategies to produce biphasic MEW scaffolds. The scaffolds were fabricated from medical-grade poly(ε-caprolactone) (PCL) to produce a five-layer 10 mm × 40 mm scaffold containing two 20 mm patterns and an interfacial region. The size of the interfacial region depended on the interfacial bonding method used: overlap was 1 mm, suture was 2 mm, and the continuous method resulted in a gradual, undefined transition region. To demonstrate the ability of the three methods to bond spatially heterogeneous scaffolds, the pore size of each pattern was varied between 1 mm and 0.5 mm. The square pattern was chosen to alternate pore sizes, while the diamond and serpentine pore sizes were kept constant at 0.5 mm and 1 mm, respectively. The pore sizes were chosen to facilitate printing and characterize the effects of the interfacial bonding methods, independent of the scaffold's suitability for tissue engineering. A control scaffold (40 mm × 10 mm single-pattern scaffold without an interfacial region) was also printed to test each pattern and pore size independently of any interfacial region. The average fiber diameter was 26.09 ± 1.90 μm. Figure 4-2 Exemplary images of each interface type are shown for square to diamond shaped scaffolds.

[0144] Next, each joining method was evaluated for its similarity to the programmed print path by comparing their morphologies using optical microscopy and scanning electron microscopy (SEM). Fiber bridging, a known defect in which fibers deviate from their intended path on adjacent fibers due to electrostatic repulsion caused by residual charge, was identified in varying amounts across all scaffolds. The overlapping scaffolds exhibited the densest interface regions, with some instances of bridging. The stitching method was more customizable, but for the chosen designs, the fiber density was slightly lower and the amount of bridging to the overlapping scaffolds was comparable. The continuous printing technique best matched the planned print path, with fewer fiber bridging defects and the lowest density. These results are consistent with the literature, as the amount of printing defects in MEW scaffolds is expected to increase in areas of higher fiber density and in areas where the jet abruptly changes direction. Therefore, as opposed to dense bands containing more fibers, the continuous interface produces a more precise imprint and exhibits a gradual transition in density between spatially heterogeneous regions.

[0145] Example 2

[0146] Effects of interface printing methods on the mechanical properties of dual-phase MEW scaffolds

[0147] Uniaxial tensile test

[0148] To evaluate the effect of the interface printing method on the mechanical properties, the biphasic MEW scaffolds were subjected to uniaxial tensile tests under a load perpendicular to the interface. Figure 5-1 Included are time lapse images taken at different stages of stretching, from left to right: unloading stage, elastic deformation stage, plastic deformation stage, and scaffold rupture. Figure 5-2 Representative stress-strain plots of continuously articulated scaffold structures are depicted, with labels corresponding to regions of i) elastic deformation, ii) plastic deformation, and iii) scaffold structure rupture. Figure 5-3 Depicted are representative stress-strain plots for the combination of the serpentine square scaffold structure with each type of interface and the control scaffold structure. Figure 5-4 Young's modulus, yield strength, and ultimate tensile strength of each type of biphasic scaffold structure with the interfacial method and the control scaffold structure are depicted (mean of n=3, error bars represent one standard deviation, significant differences were assessed using one-way ANOVA and Tukey's multiple comparison test, where statistical significance is shown on a single bar compared to every other bar on the graph, where the data had at least that level of significance, *p≤0.05, **p≤0.01, ***p≤0.001, ****p≤0.0001).

[0149] Figure 6-1 Images of a continuously bonded square-to-snake scaffold structure to which a uniaxial strain force was applied at different clamping distances are depicted. The effect of the clamping distance from the interface on the mechanical results was first investigated, and no significant differences were observed in the calculated stresses. Stress parameters including Young's modulus, yield strength, and ultimate tensile strength are depicted in Figure 6-3 、 6-4 and 6-5. exist Figure 6-3 、 6-4 In each of 6-5 and 6-5, the results depicted in order from left to right correspond to the results obtained with the clamping distances of 2.5 mm, 4.5 mm, and 6.5 mm, respectively.

[0150] However, significant variations in strain values ​​were observed, and thus relevant metrics such as yield strain could not be compared. Subsequently, the stent was clamped closer to the interface on the weaker stent side due to the limited strain testing range of the equipment. Figure 5-2A representative stress-strain diagram 501 of a continuously joined serpentine-square scaffold structure is shown. The initial elastic deformation of the scaffold structure (i) followed by plastic deformation (ii) is almost entirely explained by the weaker pattern, in this case the serpentine. The fibers of the weaker scaffold then begin to neck, resulting in strain hardening, manifested as a gradual increase in stress. When the scaffold structure reaches its ultimate tensile strength (UTS), the scaffold structure ruptures (iii), indicated by a sharp decrease in stress corresponding to the strain at which the two patterns separate from each other. Similar stress-strain curves are common in the rest of the interfacial scaffolds, such as Figure 5-3 The stress-strain diagrams and scaffold testing of the serpentine-square, square-diamond, and serpentine-diamond scaffolds show similar behavior.

[0151] The Young's modulus of the bonded scaffold is dominated by the more elastic (meaning lower modulus) of the two bonding patterns ( Figure 5-4 ). In all cases except one (continuous serpentine-square), the elasticity of the joined scaffold is equal to or greater than the elasticity of the more elastic pattern. For square-diamond and serpentine-diamond scaffolds, Young's modulus is not affected by the interface joining method. It is noteworthy that, regardless of the joining method, the joined square-diamond scaffold structure results in a scaffold structure that is significantly more elastic than any of its components. However, in the serpentine-square scaffold structure, the suture method is the only technology that leads to a more elastic scaffold structure. These results show that although most of the elastic behavior comes from the more elastic pattern, both patterns contribute to the elasticity of the final scaffold structure in some way. The yield strength indicates the stress level at which the scaffold structure will begin to deform permanently. The results show that yield is completely determined by the weaker pattern, which means the pattern with lower yield strength. Similarly, the UTS of the scaffold is mainly determined by the weaker of the two patterns, with only one exception (suture square diamond). Overall, due to the advantage of the weaker scaffold, the joining method has no significant effect on the uniaxial mechanical properties. However, other mechanical properties associated with tissue engineering scaffolds may be affected.

[0152] Flexure test

[0153] Next, we investigated how the printing method used to join two different patterns changes the flexural properties of the resulting biphasic scaffolds. To this end, we constructed Figure 7-1 Flexure testing apparatus 700 is schematically shown in FIG. Figure 7-2 is a close-up view of the dual-phase stent 701, with the imaginary dashed line 702 indicating the location of the pivot line. Scale bar = 1 mm. Figure 7-2 The test image of the multiphase scaffold with serpentine pattern is shown. The flexural stiffness data are shown in Figure 7-4(mean values ​​of n = 3, error bars indicate standard deviation, two-way ANOVA with Tukey's multiple comparison test resulted in p < 0.0001 for pattern effect and p = 0.0002 for interface effect, and no significant differences between individual means).

[0154] During the test, each stent was pivoted immediately on one side of the interface, leaving one pattern fixed and one flexed, from which the bending angle was measured and the flexural stiffness (G, measured in Nm) was subsequently calculated. The G value of the control stent was also calculated. A lower G value indicates a more flexible stent. The flexural properties depend largely on the pattern of bending the stent, such as Figure 7-4 As shown. Among the patterns tested, the serpentine scaffold structure is the most flexible (minimum flexural rigidity), followed by rhombus, 1 mm square and 0.5 mm square. Note that when the bending angle θ tends to 0 °, G tends to infinity, which means that a very hard scaffold structure that is almost not flexed has an exponentially higher G value. This is observed for both square patterns, which have significantly higher G values. It is noteworthy that the interface type significantly affects the flexural rigidity. In particular, for harder patterns, both the sutured interface and the continuous interface are able to achieve greater deflection of the harder square pattern, where the continuous interface is able to achieve the maximum increase in deflection. In contrast, overlapping interfaces generally have an adverse effect on deflection. For patterns that have been flexible, such as serpentine and rhombus, the interface has a negligible effect. Comparing the overall interface technology, a correlation can be observed, whereby the continuous interface provides the greatest degree of deflection, followed by suture and then overlap. It is noteworthy that a similar trend was observed earlier when visually analyzing the fiber density in the interface area between the various methods. Therefore, this could be a possible mechanism behind the observed bending properties whereby an interface with a higher density and more fiber fusion sites would result in greater flexural stiffness.

[0155] In summary, from the test results, in most cases, the bonded MEW scaffold will not weaken the mechanical properties of the scaffold and generally result in equal or greater elasticity compared to the individual components. In addition, when trying to achieve greater flexural properties, a continuous interface bonding approach is recommended. This is valuable in the context of interfacial tissue engineering, as specific flexural properties may be required at the interface between two regions. In the context of tissue-engineered heart valve scaffolds, the interface between the leaflets and the interleaflet triangles requires a high degree of flexibility because a large amount of cyclic bending will occur. Therefore, this strategy can continue to be utilized to further unlock the capabilities of MEW scaffolds.

[0156] Example 3

[0157] User-generated MEW scaffolds with complex continuous interfaces

[0158] Writing G-code for continuous printing of MEW scaffolds can be time consuming, especially in the case of complex scaffold architectures with multiple regions. A graphical user interface (GUI) can be written to leverage GUI technology and computing advances to more quickly generate MEW scaffolds, including customizable interfaces using continuous printing.

[0159] exist Figure 8-1 A screenshot of a GUI 800 is shown in FIG. 1 for designing a scaffold having an interface between two segments "A" and "B." The inventors previously proposed a novel MEW scaffold structure in which each segment comprises interdigitated fiber groups, wherein each group comprises fibers aligned parallel or substantially parallel to one another. This results in pores within each of the segments being of identical or substantially identical shape.

[0160] For simplicity, a GUI 800 is provided for designing a stent of the type described above. The stent further has a continuous interface as discussed herein. GUI 800 provides selectable fields 810 and 820 in which the pattern (i.e., shape) of the pores for each segment can be selected. Rectangular, serpentine, auxetic cube, and auxetic star are included as examples and are not intended to be limiting. GUI 800 also includes other fields for selecting segment length, pore size, and path type. In this example, the "Path Type" choices are "Grid," in which the fibers are laid in horizontal and vertical lines forming a rectangular grid, "Diamond," in which the fibers are laid diagonally, and "Connect," in which two or more fibers are connected. It will be appreciated that other path types can be defined and presented as options. At GUI field 830, the interface function of the continuous interface can be further set to a mathematical function by selecting one of the function options, thereby allowing the boundary between segments "A" and "B" to be indicated by a mathematical function while still being continuous. Non-limiting example functions include sine, parabola, arrow, and angled shapes. GUI field 840 also allows selection of how to vary the Y hole size between segments, including halving and fanning out, as described above with respect to FIG. 3 .

[0161] Figure 8-2 and Figure 8-3 Two exemplary MEW scaffold designs and their use are shown. Figure 8-1 In the design of the input parameters, the bracket 841 is designed using the joint and fan-out method. Figure 8-2 In the example, the stent 842 was designed to adopt a diagonally oriented path (i.e., the path type was selected as "diamond"), where fanning occurs in one section. Digital microscope images of MEW stents fabricated using PCL are shown in Figure 1. Figure 8-4 and 8-5 Shown in.

[0162] The method of the present invention has the ability to generate continuous, user-defined MEW scaffolds. This new method and its implementation using, for example, a GUI unlocks the ability to rapidly and iteratively design a variety of continuous interfacial scaffolds, providing extensive tunability of mechanical and morphological properties, which could be highly beneficial for interfacial tissue engineering applications using MEWs.

[0163] Example 4

[0164] Morphological Analysis of the Aortic Heart Valve Interface

[0165] In the experiments conducted, multimodal imaging was used to examine the morphology of the aortic heart valve interface region as a basis for establishing a functional, bioinspired design of the MEW-scaffold interface. Subsequently, three methods for joining heterogeneous scaffolds were investigated to understand their effects on tensile and bending properties. Combining investigations of human tissue with research on interfacing methods, a complete biomimetic heart valve interface was designed, fabricated, and tested.

[0166] The aortic valve leaflets are responsible for most of the valve's overall function and have been the focus of previous research. Additionally, other regions of the aortic valve are known to contribute significantly to the valve's function. These regions have been categorized in different ways. For the purposes of this specification, Figure 9-1 As shown, these regions are divided into the leaflets, sinuses of Valsalva, annulus, commissures, and interleaflet triangles. Each region has different concentrations and orientations of extracellular matrix (ECM) components, including collagen and elastin fibers, which are of particular interest when considering the mechanical function of the valve. Collagen fibers are the primary load-bearing component of the ECM and are responsible for the anisotropic mechanical properties and J-shaped stress-strain response of the tissue. Elastin fibers play a role in the low-strain performance of the valve, but more importantly regulate the orientation of the collagen fibers, ensuring that they return to their preloaded state between cycles. Therefore, while all components of the ECM play an important role in valve function, our studies focused on the orientation of the collagen fibers to inspire the MEW stent design. Therefore, it is important to understand what is known about the orientation of collagen fibers in the heterogeneous regions of the aortic valve, which is briefly summarized below.

[0167] The aortic valve leaflets contain three distinct layers: the fibrous membrane, the spongiosa, and the ventricular membrane (from the aorta to the ventricle). The tensile load-bearing layers are the fibrous membrane and the ventricular membrane, which are primarily composed of circumferentially and radially arranged collagen fibers, respectively. The sinuses of Valsalva are similar in composition to the aortic wall and consist of three distinct layers, of which the inner (intima) and outer (adventitia) layers are primarily composed of longitudinally arranged collagen fibers, while the middle (media) layer is primarily composed of circumferentially arranged fibers. The annulus is a fibrous structure that connects the leaflets and the sinus wall to the left ventricle and is primarily composed of high-density, circumferentially arranged collagen bundles, resulting in relatively rigid mechanical properties. The commissures are present at the point where the free edges of the two leaflets meet. By helping to transmit forces between the leaflets and the surrounding aortic root, the commissures play a vital role in supporting the bending properties required to open the valve during diastole and to withstand the high tensile loads when the valve closes during systole

[66] . In terms of microstructure, collagen fibers extend primarily radially from the leaflets, wind through the commissure region and anchor into the aortic wall, enabling force transmission between the leaflets and the root. The interleaflet triangle is dominated by the regions below the commissure, between the leaflets and above the annulus. Within the interleaflet triangle, the orientation of the fiber microstructure is largely unknown; however, it is thought to be primarily in the direction of the annulus, i.e., circumferential. Notably, while the collagen microstructure of these regions is known, the specific nature of how the collagen fibers are oriented in the interface between these regions remains to be investigated. This could provide valuable insights to inform scaffold designs that strive to mimic natural fiber morphology. Therefore, the inventors began the problem-solving process by investigating whether high-resolution multimodal imaging of porcine tissue could provide insight into some of these unknowns.

[0168] Experiments and analyses using porcine tissue have been widely used in cardiovascular research due to the greater availability of tissue that shares anatomical and hemodynamic similarities with humans. When analyzing collagen microstructure in these studies, local fiber orientations were consistent between human and porcine species. Consequently, porcine tissue has been used in biomechanical studies and to elucidate fiber alignment using microscopy.

[0169] like Figure 10 As schematically depicted in Figure 1, porcine aortic valve tissue 1001 was fixed using a hybrid immersion fixation and hydrostatic expansion method to maintain near-physiological diastolic conditions. First, micro-computed tomography (micro-CT) was used for 3D reconstruction of the valve structure to facilitate macroscopic identification of regions of interest. Next, the tissue was longitudinally sectioned through the wall of the aortic root, parallel to the direction of blood flow, with a thickness of 250 μm for further analysis by second harmonic generation (SHG) imaging. Figure 9-2Depicted are a 3D reconstructed valve architecture 901 and slices 902 viewed from the i) coronal (external) direction, ii) sagittal direction, iii) axial direction, and iv) coronal (internal) direction. This slice orientation allows imaging of the commissures and interleaflet triangles, the interfaces between them, and the interfaces with the leaflets, sinuses, and annulus. Collecting three slices into the aortic wall enables confirmation of orientation information in the third dimension. SHG is an optical microscopy technique that is ideally suited for imaging collagen from the tissue scale to the molecular scale due to the second-order nonlinear properties of collagen fibers.

[0170] In SHG images (e.g. Figure 9-3 ) to identify the microstructure of a specific heart valve region, including the commissures of two adjacent leaflets and the fibrous and ventricular layers. In the commissures, a large number of collagen fibers are densely interwoven and tangled between vertical and horizontal directions. Three different groups of collagen fibers travel from the top of the commissure down into the leaflet. The left and right sides of these groups merge into the fibrous layers of the two adjacent leaflets, which can be identified by their distinct rounded structure, with collagen fibers traveling in and out of the leaflet, as shown by the weakened SHG signal. Remarkably, we observed that the fibers of the central group travel down from the top of the commissure for approximately 500 μm before splitting into the ventricular layer of each leaflet, which can be identified by the vertical alignment of collagen. To our knowledge, the orientation and extent of the interface between the commissure and the fibrous and ventricular layers of the two adjacent leaflets has been unknown so far. In deeper images of the commissures, we observed highly aligned collagen fibers traveling in a horizontal direction from the sinus and then coalescing into larger bundles as they turn to travel vertically downward, at which point the fibers are strongly aligned and tightly bundled again ( Figure 9-4 ).

[0171] SHG imaging of the interlobular triangle reveals horizontally oriented collagen fiber bundles emerging from the leaflets and the sinuses on both sides (see Figure 9-5 Fibers are regularly intertwined as they transition from horizontal to diagonal and then primarily vertical orientation. The vertically aligned bundles that exit the top of the image continue into the commissure. The change in fiber orientation is gradual, with the fibers gathering into discrete bundles as they rotate before achieving vertical or diagonal alignment.

[0172] To validate the SHG collagen fiber orientation data, correlated focused ion beam scanning electron microscopy (FIBSEM) imaging was performed on the same sample. A specific region from the SHG image was selected and co-registered to the FIBSEM slice ( Figure 11-1 At low magnification, FIBSEM sections show that collagen fiber bundles fill the spaces between valve fiber cells ( Figure 11-2 ). Figure 11-3 yes Figure 11-2A higher magnification image of the framed area 1101 shown in FIG. At higher magnification, it can be seen that the individual collagen fiber cross sections are aggregated into bundles, running primarily in the circumferential direction C, and one bundle changes from radial to longitudinal (R, L). This is consistent with the Figure 11-1 The results are consistent with those observed from SHG imaging in the 3D images, in which collagen fibers mainly travel in a circumferential direction perpendicular to the plane of the slice, while gradually turning to a longitudinal motion parallel to the slice plane.

[0173] Based on multimodal imaging analysis of porcine aortic valves, we can summarize the following: 1) the commissures consist of a complex interwoven network of collagen fibers originating from multiple directions, which then coalesce into a primarily perpendicular alignment at the core of the commissures; 2) within the interleaflet triangular region, collagen fiber bundles exhibit a regular diagonal weave before coalescing toward a predominantly perpendicular alignment, with more bundles forming as they transition to the commissures; and 3) fiber orientation changes at the interface are gradual and generally manifest as a transition from uniformly aligned, large sheets of fibers to more discrete bundles as the fibers rotate at increasingly acute angles. Due to the complexity of developing a sample preparation method compatible with accurate cross-modality correlation imaging and the considerable time required to perform this method, this study was limited to n = 1 sample. Observations from high-resolution multimodal imaging analysis of the aortic heart valve interface were used in this study as inspiration for the design of an interfacial fiber heart valve scaffold.

[0174] Tissue source and preparation

[0175] The collection and use of porcine animal tissue in this work was approved by the University of Western Australia Biosafety Committee (F 69199). Porcine hearts were obtained from the University of Western Australia Large Animal Laboratory and tissue was removed within 2 hours of euthanasia. Following removal, the hearts were immediately dissected, leaving the aortic root and its valves, sinuses, the first few millimeters of the coronary arteries, and a portion of the ascending aorta intact.

[0176] Primary fixation of tissue (n = 1) was performed in fresh 4% paraformaldehyde solution (Cat#C007, ProSciTech, Australia) prepared in 0.1 M phosphate buffer (PB), pH 7.4. Fixation was performed for 1 hour at room temperature and at a diastolic pressure equivalent to 80 mmHg using a hybrid immersion fixation and hydrostatic expansion device as shown in Figure S1 (Supporting Information). Secondary fixation was performed by immersion overnight in 2.5% glutaraldehyde (Cat. No. EMS 16400; ProSciTech, Australia) in 0.1 M PB, pH 7.4. Excess tissue was removed by dissection and then stored in glutaraldehyde fixative solution at 4°C.

[0177] Microcomputed tomography

[0178] During imaging, tissue samples were kept moist under a paper towel soaked in glutaraldehyde solution and placed in a sealed zip-lock bag. MicroCT imaging (Skyscan 1176, Bruker-microCT, Kontich, Belgium) was performed at a source voltage of 45 kV, a source current of 556 μA, an exposure time of 86 ms, and a resolution of 34.81 μm / pixel, using a 0.2 mm aluminum filter, a 360° rotation of 0.7°, and 2 / image frame averaging. MicroCT data were segmented and reconstructed using in-house software before being imported into STAR-CCM+ (v16.04.012-R8, Siemens) for further smoothing and processing.

[0179] Second harmonic imaging

[0180] 5% agarose (Agarose LE, analytical grade, Promega, Australia) embedding solution was used at 65°C. The tissue sample was removed from the fixative solution, blotted dry, and then placed in a custom-sized (approximately 30 mm cubed) Lego (The Lego Group, Denmark) mold that contains the entire volume of the tissue and minimizes the amount of agarose solution used. The agarose solution was poured around the tissue, ensuring that minimal bubbles were left. The embedded tissue was cooled to room temperature and then stored at 4°C overnight. For slicing, the solid gel / tissue block was removed from the mold and super-glued to a vibrating microtome table (vibrating microtome 3000, The Vibratome Company, St. Louis, MO, USA). After slicing the block to the area of ​​interest, three serial sections were taken at a thickness of 250 μm. Optical imaging was then performed on these sections using an inverted A1RMP multiphoton microscope (Nikon) equipped with a 10× / 0.45NA objective (Nikon) and a tunable laser (10 mW output, 900 nm emission wavelength). All images were captured using NIS-Elements AR software (v5.30.02; Nikon).

[0181] Electron microscopy of tissue

[0182] Electron microscopy samples were prepared from sections after SHG imaging using a BioWave Pro microwave system (Pelco) by microwave-assisted processing. Briefly, samples were osmated using the R-OTO method, stained en bloc with aqueous uranyl acetate and lead aspartate solutions, and then dehydrated through a graded ethanol series (80%, 90%, 95%, 100%, 100% (v / v)) and propylene oxide (100%, 100% (v / v)). Sample infiltration with Araldite 502 / Embed 812 was performed in propylene oxide under vacuum (25%, 50%, 75%, 100%, 100% (v / v)) via a graded concentration series. Samples were polymerized at 60°C for 48 hours and trimmed on an ultramicrotome (Leica UC6, Leica Biosystems) in preparation for imaging on a FEI Helios Nanolab G3CX DualBeam FIB-SEM (Thermo Fischer Scientific). The target area was milled using a gallium FIB current of 65 nA at 30 kV, and the block face was then imaged with backscattered electrons at an accelerating voltage of 2 kV in magnetic immersion mode.

[0183] Example 5

[0184] Design, fabrication, and testing of a bioinspired aortic valve interface stent

[0185] Leveraging new microstructural information of the aortic valve interface obtained from multimodal imaging studies, a bioinspired MEW scaffold was designed using a complex continuous interface in the aortic heart valve interface region. Figure 9-5 Orientational analysis of collagen structures enables a quantitative understanding of microfibrillar collagen orientation. For example, see Figure 12-1 、 12-2 Figures 12-3 and 12-3 show color-mapped orientation analysis of SHG images of interlobular triangular regions (each scale bar = 0.5 mm). In the tissue sample, the fiber distribution was observed to gradually change from 0° (horizontal) to a peak at ±50° (diagonal), and then gradually change again to a ±90° (vertical) fiber alignment. Notably, few fibers were completely vertically oriented, with most fibers woven in a regular diagonal pattern with a slight vertical tilt.

[0186] The intricately woven fibril structure of the interlobular triangular region is simplified into schematic form ( Figure 12-2) to facilitate stent design. The pattern in the leaflet region is based on the heart valve leaflet design previously published by the inventors (Saidy et al., (2019) Small 15(24), https: / / doi.org / 10.1002 / smll.201900873). For the interleaflet region 1201, instead of being discretized into commissures, interleaflet triangles, and annulus, a gradient diamond pattern is used, such as Figure 12-3 The gradient transitions from a predominantly horizontal alignment at the bottom, which is intended to mimic the circumferential fibers present in the annulus, to a predominantly vertical alignment at the top, which mimics the longitudinal fibers present at the commissures (Fig. Figure 9-3 、 9-4 (see in the ).

[0187] The fiber design avoids completely horizontal or vertical orientation because the straight new fiber pattern is more rigid and weaker than a diamond pattern of the same porosity, such as from Figure 5-4 This can be seen from the data shown in . When comparing the fibers in the MEW scaffold design with those previously analyzed in native tissue, similar fiber orientation distributions were achieved, as shown in Figure 12-4 The resulting scaffold design thus combines natural tissue observations with continuous interfaces to produce a biomimetic scaffold design with gradient porosity, region-specific layer numbers, tailored fiber orientation, and spatially heterogeneous regions.

[0188] The above-mentioned method for designing MEW heart valve stents can be extended to design MEW stents for other soft tissues.

[0189] Figure 12-5 The scaffold 1201 of the configuration shown in FIG was successfully fabricated from PCL with an average fiber diameter of 27.69 ± 4.66 μm. Figure 12-6 Snapshots of the deposited fiber taken at times 1 second, 2 seconds, 6 seconds, 29 seconds, and 74 seconds after the start of MEW printing are shown to illustrate the continuous printing path.

[0190] The continuous printing path was designed to employ a joining technique that continuously transitioned from 5 layers in the leaflet region to 10 layers in the rest of the scaffold. In some cases, where double print layers were deposited due to the joining method, fiber fusion occurred simultaneously, resulting in thicker fibers instead of two layers of fibers. Figure 12-7Fused, thicker fibers 1202 can be seen. The average fiber diameters of the unfused and fused fibers were 24.34 ± 0.75 μm and 33.27 ± 2.30 μm, respectively. Comparison of the cross-sectional area of ​​the fused fibers with the unfused fibers confirmed the fiber fusion phenomenon, with the fused fibers having approximately twice (187 ± 14%) the area. This is likely due to the very short time between fiber depositions on the same path, causing the deposited fibers to remain at a high temperature. This is reminiscent of the morphological arrangement of collagen fibers found in natural tissue, where smaller collagen fiber bundles connect together to form larger, tightly packed fiber bundles as they change orientation.

[0191] Then as Figure 13-1 As shown, the scaffold was mechanically characterized under biaxial physiologically relevant strains in the "circumferential" and longitudinal directions. Tissue-like characteristics such as anisotropy and viscoelasticity were investigated using relative strain-related metrics to understand them. Furthermore, it is important to note that this scaffold contains leaflets, interleaflet triangles, annuli, and commissural regions, meaning that comparisons with region-specific data are not strictly physiologically accurate. Attempts were made to evaluate similar loading regimes in porcine tissue, however, due to the difficulty in clamping the tissue to the mechanical tester and the variable cross-sectional area, no reliable quantifiable data was available for comparison. Notably, it was observed that in all porcine tissue samples (n=12), failure never occurred in the interface region. Therefore, comparisons were made between the scaffold and reported native tissue values ​​in regions that were as similar as possible to the scaffold. In future studies, hemodynamic testing will ensure normal physiological loading of such scaffolds.

[0192] The ratio of strain rates in the circumferential to longitudinal directions was chosen to be approximately 2:1. When tested to failure, the scaffold exhibited similar Young's moduli in both directions, while the yield strength in the circumferential direction was approximately twice that of the longitudinal direction (see Figure 13-2 ). The strength of the stent can be improved by adjusting the number of layers or fiber diameter. The strain-displacement vector plot shows a complex regional anisotropic deformation behavior, in which the circumferential strain and longitudinal strain are mainly explained by the leaflets and interleaflet triangular regions, respectively. However, the strain-displacement vectors show a gradual change in direction and magnitude between these regions, indicating a smooth transition of loads across the interface. It is worth noting that the yield strains in the circumferential and longitudinal directions are 27.7±6.9% and 16.6±5.4%, respectively, both of which are higher than the average maximum strains reported in the literature on different regions of the aortic valve. This means that the stent will remain within the elastic region and will not deform plastically under physiological strains.

[0193] The hysteresis of the stent structure was then determined by cyclically testing the samples at constant strains of 20% and 10% in the circumferential and radial directions, respectively, and calculating the ratio of the area under the unloading curve to the area under the loading curve. Figure 13-3 Figure 3 shows a circumferential strain 1301 and a longitudinal strain 1302. A constant low level of hysteresis (approximately 20% energy loss per cycle) was observed in both directions. Our data are consistent with the 17% hysteresis reported in the literature for porcine aortic heart valves. The relaxation behavior of the stent was characterized using 5% and 2.5% strain increments in the circumferential and radial directions, respectively. Figure 13-6 Then, after each strain step, the scaffold was allowed to relax for 1000 s, which is a sufficient time frame to exhibit most of the relaxation behavior. Figure 13-5 Stress-time plots representing relaxation behavior are shown. At all strain levels, the stent exhibited rapid initial relaxation before stabilizing to an asymptote, which was then recorded as a percentage of relaxation. A consistent degree of relaxation of 27.6 ± 1.5% was maintained in the circumferential direction, while the radial direction exhibited an initial large relaxation before stabilizing to 30.6 ± 1.7%. These values ​​are very similar to those reported for porcine aortic heart valve relaxation. In summary, the bioinspired aortic heart valve interface stent thus exhibits characteristics reminiscent of natural tissue, including anisotropy, hysteresis, and yield strain within natural tissue.

[0194] In this paper, a continuous interface across the structural regions of a MEW scaffold is provided. Complex functions are used to define the printing paths between discrete endpoints in the structural regions connected by the interface regions. Consequently, fibers adopt complex shapes as they transition from one region through the interface region to the next, which is the first region in the structure. A simplified representation of the microstructural organization of collagen fibers in the interface regions of biological tissues can be exploited when designing MEW tissue scaffolds. For example, during testing and analysis, high-resolution correlative multimodal imaging was used to obtain the microstructural organization of collagen fibers in the interface region of an aortic heart valve. This was used to inspire the functional design of MEW heart valve scaffolds with biomimetic interface regions, including for applications in heart valve tissue engineering. Systematic morphological and mechanical studies of three different methods of joining biphasic MEW scaffolds revealed that the weaker regions dominated the tensile mechanical response. Regarding flexural stiffness, a property often overlooked in scaffold design, the continuously joined biphasic scaffolds demonstrated enhanced flexibility of the resulting constructs. Crucially, a novel software and accompanying GUI not only enabled the rapid design and G-code generation of a range of MEW composite scaffold designs, but also generated the continuous interface boundaries connecting these regions. This software unlocks new capabilities for MEW in the field of interfacial tissue engineering. Subsequently, we designed, fabricated, and tested a novel bioinspired scaffold for the aortic heart valve interface region, demonstrating for the first time a scaffold incorporating a continuous interface, gradient porosity, region-specific layer numbers, and customized fiber orientation within a single biomimetic design. When tested under physiologically relevant biaxial conditions, the scaffold exhibited promising tissue-like behaviors, including strain yielding, hysteresis, and relaxation, similar to native tissue.

[0195] MEW bracket manufacturing

[0196] The scaffolds were fabricated from medical grade poly(ε-caprolactone) (PCL) using an in-house constructed MEW apparatus as previously reported. The specific processing parameters used were: extrusion driven at 100 kPa; through a 23G metal needle; a working distance of 3 mm; a translation speed of 400 mm / min; a voltage of 4.4 ± 0.1 kV applied to the spinneret and the grounded collector plate; with an 86°C ring heater and a 31.5°C bed heater. The exact processing parameters used varied slightly depending on the ambient conditions of the day. All biphasic scaffolds used for tensile and flexural testing had dimensions of 10 mm high by 40 mm wide (20 mm wide per pattern). It was found that this fabrication setup allowed for fiber fusion within the same layer as well as adequate bonding between different layers. In some cases, complete fusion between layers was not required where the layers did not have the same configuration, or was undesirable depending on the degree of flexibility required for the overall layered structure.

[0197] MEW stent imaging and measurement

[0198] The MEW scaffolds were first imaged using a Hirox RH-2000 digital microscope (Hirox, Europe) at a resolution of 4.51 μm / pixel. The accompanying software was used to measure the fiber diameter at three locations for n = 3 scaffolds of each type. The scaffolds were then cut using scissors and secured to SEM stubs using double-sided carbon tape. The stubs were then sputter-coated with gold, two at a time, over 60 seconds using a JEOL Smart Coater (JEOL, USA). SEM images were then acquired using a JCM-6000 desktop SEM (JEOL, USA) at an accelerating voltage of 10 kV, 30× magnification, high probe current, and high vacuum mode.

[0199] Uniaxial tensile test

[0200] The scaffolds were subjected to uniaxial mechanical tensile testing across the interface using a CellScale Biotester (CellScale, University of Waterloo, Canada) equipped with a 1.5 N load cell. The samples (n = 3 for each scaffold type) were fixed with custom 3D printed micro-fixtures and suspended in air at room temperature. A displacement rate of 1% length / min was used for uniaxial testing. The resulting force-displacement data were used to plot stress-strain curves, where strain is defined as engineering strain (change in length / initial length). For all calculations, the initial length was kept constant at 15 mm. The cross-sectional area was quantified as the height of the scaffold wall, rather than the height of the peak. Six measurements were performed on n = 3 samples of each biphasic scaffold. The cross-sectional area was observed to vary within a single scaffold between the two patterns and the interface. However, since most of the deformation occurred in the weaker pattern, as confirmed by visual analysis, the cross-sectional area of ​​the weaker pattern was used to calculate the associated stress. Average value per pattern: Serpentine = 1.06 ± 0.03 mm 2 , rhombus = 1.25 ± 0.20 mm 2 , 1 mm square = 1.40 ± 0.21 mm 2 , 0.5 mm square = 1.16 ± 0.04 mm 2 Young's modulus was calculated from the slope of the stress-strain curve in the steepest linear region, with the strain range depending on the sample / pattern. Yield strength was calculated as the plateau point of the stress-strain curve. Ultimate tensile strength was determined as the maximum stress reached for each sample.

[0201] Biaxial tensile testing

[0202] Biaxial tensile testing was performed using the same equipment as the uniaxial test with a 5N load cell. A preload of 50mN was applied before each test to ensure that the initial state of the stent was consistent. Young's modulus and yield strength were calculated using the same scheme as the uniaxial test. Cyclic testing was performed for 1 preload cycle (to establish the in vivo state of the material) and 9 subsequent loading cycles, which were then plotted as stress-strain. For each cycle, hysteresis was calculated as the ratio of the area under the unloading curve to the area under the loading curve. 5% and 2.5% strain increments were used in the circumferential and radial directions, respectively, to characterize the relaxation behavior of the stent structure (Figure 8D). The stent structure was then relaxed for 1000 seconds, which is enough time to show most of the relaxation behavior. The difference between the maximum stress and the minimum stress in each step is the relaxation percentage.

[0203] Bending test

[0204] Flexural testing was performed using the Peirce cantilever test method (ASTM D1388). A custom test setup was designed in Fusion 360 (Autodesk, USA) and printed using Prusament PLA (Prusa Research, Czech Republic) and a Prusa MK3s+ (Prusa Research, Czech Republic). Six independent angle measurements were taken for each flexural pattern, three times in one direction and three more times with the scaffold flipped to account for the natural warping of the scaffold after printing.

[0205] Complex G-code generator GUI and path visualization

[0206] An in-house program was developed using Python (Python Software Foundation, USA) to integrate continuous interface design into spatially heterogeneous G-code for MEW printing. A script was created to generate a GUI, into which the desired parameters for the bracket could be entered. Once entered, the interface executed additional scripts that performed the necessary calculations to plot the data for the desired print path. This data was then converted by another script into a G-code notepad file for input into the MEW printer software. To visualize the G-code, additional custom Python-based software was built that imported the raw G-code and produced static and dynamic renderings of the print path, with customizable scale, color, line width, and path speed.

[0207] Directional Mapping

[0208] Orientation analysis of the SHG images and the new scaffolds was performed using the plugin OrientationJ

[90] for the software Fiji (National Institutes of Health, USA).

[91] Using the OrientationJ analysis module, a hue-saturation-brightness color map was calculated using 2-pixel cube ridges. The orientation distribution map was then plotted using the OrientationJ distribution module.

[0209] Statistical analysis

[0210] Unless otherwise stated, all data are presented as mean ± standard deviation of n = 3 samples. For uniaxial tensile test data, significant differences were assessed using one-way ANOVA with Tukey's multiple comparison test. For flexural stiffness data, significant differences were assessed using two-way ANOVA with Tukey's multiple comparison test. For biaxial tensile test data, significant differences were assessed using parametric, ratio paired t-test. The α value for all tests was 0.05. All statistical analyses were performed using GraphPad Prism software (GraphPad, San Diego, USA).

[0211] As mentioned above, MEW printing contrasts with conventional MEW printing, in which layers comprising separate structural regions are not printed in a single, continuous run. Instead, regions of the layer are printed separately and then stitched together to form the layer. There is no perceived mechanical advantage to continuous printing across the entire region. The contrast with conventional thinking is even greater in embodiments where fibers are intentionally converged into thicker fibers. This is considered "fiber bridging," and conventional wisdom considers it a drawback that should be avoided. This is because in MEW, the ability to reduce pore size is limited due to electrostatic fiber repulsion, which is inherent to the MEW printing process. In MEW printing, high voltage electrostatically charges the molten jet, and as the molten jet cools, the charge is trapped. Those skilled in the art understand this phenomenon and would be expected to consider fiber convergence as something to be generally avoided. However, in some embodiments of the present method, "fiber bridging" is intentionally induced in a controlled manner. The MEW printing parameters presented above allow for controlled bonding between two merging fibers to form a single, thicker fiber. This opens up the possibility of creating biomimetic designs for MEW scaffolds, leveraging data on fiber distribution that can be obtained using imaging modalities. Scaffolds designed using this method can be used to create soft tissue implants with biodegradable materials or to create non-biodegradable implantable devices.

[0212] Example 6

[0213] Gradient porosity scaffolds

[0214] G-code is the language used by most 3D printers, specifying Cartesian (xyz) coordinates that the printer moves to in a specific order. A G-code generator was created using Python, allowing users to input desired design parameters and produce output G-code to control the MEW printer, along with an image preview of the scaffold architecture. A novel feature of this generator is the ability to specify a gradient of porosity.

[0215] Gradients are defined mathematically and can be customized ( Figure 16 ). The gradient function is a linear function determined by the following formula:

[0216]

[0217] where p n is the pore size (any positive number in mm) and m is the gradient coefficient (any non-zero number where a positive m will increase the pore size and a negative m will decrease the pore size).

[0218] Since each pore has two dimensions (length (X) and width (Y)), a gradient can be applied to one or both dimensions. These are called X gradients or Y gradients, respectively, indicating that the pore size is changing.

[0219] The entire scaffold structure itself is also two-dimensional, so the pore size can vary in one or both of these directions (horizontally or vertically). For simplicity, in this example, the gradient is only applied in the horizontal direction.

[0220] This gives two types of gradients:

[0221] An X gradient applied horizontally. This can be viewed as changing the aperture parallel to the gradient direction, and is therefore called parallel gradient .

[0222] A Y gradient applied horizontally. This can be viewed as changing the aperture orthogonal to the gradient direction, and is therefore called Orthogonal gradient .

[0223] Here, the gradient is applied to three different patterns / shapes (rectangular, diamond, or serpentine). Note that in the case of the serpentine stent, the wavelength, amplitude, or both can be varied.

[0224] A gradient is a gradual spatial variation in a property (geometric, mechanical, compositional, chemical or more). The scaffolds of the present invention produce multiple gradient properties within a single scaffold. This is already apparent geometrically ( Figure 16-19 ), however, there are also gradient mechanical properties ( Figure 20 ). Note that although not shown here, other properties such as gradient tissue / cellular composition can potentially be supported / achieved by using these scaffolds.

[0225] To understand the effect of the gradient on mechanical properties, constrained biaxial tensile testing was performed. This is where the scaffold is clamped on its four edges and then pulled in the horizontal direction while being held (constrained) in the vertical direction. This allows for observation of local variations in strain, the amount of deformation that occurs in an object, throughout the scaffold. Figure 20 Figure 3 shows the highly heterogeneous mechanical properties produced by a gradient scaffold structure. Note that a completely uniform / homogeneous scaffold design would always display the same color. The novelty here is the ability to generate a gradient in mechanical properties throughout a single scaffold. This invention provides the ability to generate gradients not only in one direction but also in complex arrangements, which could be beneficial for tissue-related applications.

[0226] Variations and modifications may be made to the portions previously described without departing from the spirit or scope of the present disclosure.

[0227] The matters set forth in the foregoing description and accompanying drawings are provided by way of illustration only and not as a limitation. While particular embodiments have been shown and described, it will be apparent to those skilled in the art that changes and modifications may be made without departing from the broader aspects of the present invention's contribution. The actual scope of the protection sought is intended to be defined in the appended claims when viewed in their proper perspective based on the prior art.

[0228] In the following claims and the foregoing description of the invention, unless the context requires otherwise due to explicit language or necessary meaning, the word "comprise" or variations such as "include" or "comprising" are used in an inclusive sense, i.e., specifying the presence of stated features but not excluding the presence or addition of other features in various embodiments of the invention.

Claims

1. A molten electrowriting soft tissue scaffold, comprising: a first region having one or more groups of fibers and a second region having one or more groups of fibers; and An interface region joining the first region and the second region is electrowritten along with the first region and the second region in a continuous print path such that fibers in the interface region are each joined between corresponding pairs of fibers in the first region and the second region.

2. The bracket according to claim 1, wherein The path of at least one of the fibers within the interface region is defined manually or by a mathematical function.

3. The bracket according to claim 2, wherein: The function is defined such that the fiber has a complex shape as the fiber transitions from one of the first region or the second region through the interface region to the other of the first region or the second region.

4. A stent according to any one of the preceding claims, wherein The second region has a higher porosity than the first region, or vice versa.

5. The bracket according to claim 4, wherein The higher porosity in the second region is formed at least in part by connecting every two or more adjacent fibers in the first region as they transition into the second region.

6. The bracket according to claim 4 or claim 5, wherein: The higher porosity in the second region is formed at least in part by causing the fibers to fan outward as they transition into the second region.

7. The stent according to any one of claims 4 to 6, wherein: The higher porosity in the second region is formed at least in part by arranging the continuous printing path so that another set of fibers in the first region are deposited offset from the other fibers by a distance less than the pore size.

8. A stent according to any one of the preceding claims, wherein The first region or the second region or both comprise a first set of fibers arranged generally parallel to one another and a second set of fibers arranged generally parallel to one another, the second set of fibers being arranged at an angle and preferably transversely relative to the first set of fibers, each fiber in the second set of fibers having a serpentine arrangement having defined valleys and peaks.

9. A stent according to any one of the preceding claims, wherein: The first region and the second region are heterogeneous in that they differ from each other in one or more spatial parameters.

10. A stent according to any preceding claim comprising multiple layers of fibres. The stent of claim 10 , having portions having different numbers of layers than different portions.

12. A stent according to any one of the preceding claims, which is a heart valve stent.

13. The bracket according to claim 12, wherein: The first region is a leaflet, and the second region is an interlobular triangle.

14. A method for providing a melt electrowriting soft tissue scaffold having at least two structurally heterogeneous regions and an interface region therebetween, the method comprising providing a printing path along which a polymer melt material is continuously extruded during a melt electrowriting process to form the heterogeneous structural region and the interface region in one printing run.

15. A method according to claim 14, comprising causing the continuous printing path to be defined in segments such that it is defined by different functions in the heterogeneous region and in the interface region.

16. The method according to claim 15, wherein The one or more functions defining the continuous print path are mathematical functions that provide a curved or serpentine shape as the continuous print path traverses the interface region.

17. The scaffold of any one of claims 1 to 12, comprising a gradient porosity, wherein the size of the pores gradually increases or decreases spatially across the scaffold in one or both directions.

18. The bracket according to claim 17, wherein The shape of the holes in the bracket is selected from: rectangle, diamond or serpentine.

19. The stent according to any one of claims 1 to 12, wherein The scaffold comprises multiple layers of fibers.

20. The stent according to claim 19, wherein The multi-layered fibers comprise gradients in geometrical features and / or mechanical properties.

21. The stent according to any one of claims 17 to 20, wherein The scaffold comprising gradient porosity is used to manufacture heart valves.

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