Manufacture of Biomimetic Auricular Surgical Implants Using 3D Printed High Density Polyethylene Microfibers

This study demonstrates a new approach to manufacturing biomimetic auricular surgical implants using melt electrowriting (MEW) technology to fabricate microfiber high‐density polyethylene (HDPE) scaffolds. An emerging filament‐driven printhead and MEW printer, termed the “MEWron”, is used to enable precise control over the material extrusion process and fiber formation. By predicting the optimal extrusion conditions, continuous and uninterrupted fiber production is achieved, enabling further optimization of filament‐driven MEW fibers with a diameter of 60.5 ± 2.6 µm. As a case study, an application of microfiber HDPE fabrication is selected that comprised the design and fabrication of personalized auricular (ear) surgical implants, specifically tailored to match the unique morphology of individual patients. Patient‐specific implant models matched to the natural shape and structure of the human ear are successfully fabricated. Furthermore, the manufactured implants exhibit excellent mechanical properties, offering a 13‐fold increase in tensile stiffness compared to MEW PCL scaffolds. Overall, this research demonstrates the feasibility and potential of MEW‐based HDPE implants as a promising alternative to traditional auricular reconstruction methods, offering an alternative avenue for improved patient outcomes and enhanced aesthetic results.


Introduction
Microtia is a congenital condition characterized by the underdevelopment or absence of the outer ear, also known as the DOI: 10.1002/admt.202301190auricle. [1]Surgical interventions for auricular reconstruction for patients with microtia or traumatic injury to the external ear remain a complex clinical challenge.In addition to microtia, there is a significant incidence of facial injuries that require reconstruction of the outer ear.These injuries can result from accidents, trauma, burns, or congenital malformations.Reconstruction of the outer ear is crucial for functional, aesthetic and psychological reasons, as the ear plays a vital role in sound localization and facial symmetry. [2]There are a number of surgical options for auricular reconstruction, each with their own benefits, risks and complications.Common options include the use of autologous rib cartilage, porous surgical implants and osseointegrated prostheses. [3]Developing innovative and personalized approaches for auricular reconstruction holds great importance in improving the quality of life for individuals with microtia and those who require ear reconstruction due to facial injuries.Complication rates using existing surgical implants remain high, in part due to the significant mismatch in mechanical stiffness between the high strength polymer surgical implants and the soft tissue and skin used to complete the reconstruction.Porous high density polyethylene (HDPE) surgical implants, specifically MEDPOR (Stryker, USA), have emerged as the clinical gold standard for implant-based reconstruction in individuals with microtia or those requiring outer ear reconstruction. [4]orous HDPE implants offer several advantages, including biocompatibility, durability, and the ability to be customized to match the patient's anatomy.However, despite their widespread use, complications of existing porous HDPE implants are reported in up to 12% of cases, [4] with infection a concern, as well as protrusion through the skin, known as extrusion.This can occur due to inadequate tissue coverage, improper implant placement or simply accidental damage to the implanted structure.The reported rates of extrusion vary but have been documented to be ≈4-5%. [3,5]Continued research and advancements in implant design and surgical protocols are essential for further reducing the occurrence of these complications, minimizing the need for multiple surgeries, and optimizing the outcomes of auricular reconstruction procedures.
3D printing has emerged as a valuable tool in addressing both the need for patient specificity as well as potentially minimizing the risk of complications in auricular reconstruction. [6]By utilizing one of several 3D printing technologies available, it is now possible to create patient-specific implants that closely match an individual's auricular morphology, leading to improved aesthetic outcomes. [7]The ability to precisely replicate the intricate structures and contours of the outer ear allows for a more natural and harmonious appearance post-surgery.Furthermore, 3D printing enables the design and fabrication of implants with biomimetic mechanical properties. [8]By tailoring the porosity and structural characteristics of the surgical implant, it is possible to locally control the regionally specific mechanical properties to mimic native cartilage. [9]The goal of this approach is to minimize the risk of complications such as infection, extrusion, or implant failure through the use of an emerging 3D printing technology that produces finer structures.The combination of patient-specific customization and achieving complex mechanical properties offered by 3D printing holds significant promise in advancing the field of auricular reconstruction, leading to enhanced patient outcomes and reduced complications.
To realize this biomimetic approach, material selection plays a critical role for ensuring long-term shape fidelity and aesthetic outcomes.Since the vast majority of routinely implanted auricular surgical implants are manufactured from porous HDPE that does not significantly degrade over the lifetime of the patient, it is known that using such thermoplastics offers a permanent mechanical support construct. [10]However, 3D printing approaches using biocompatible thermoplastics have largely focused on the use of poly(-caprolactone) (PCL), [9,11,12] a lowmelting temperature polymer with a degradation profile ranging from several months to several years. [13]PCL has been extensively used throughout tissue engineering research, using many different fabrication technologies. [14,15]While degradable materials offer the promise of fully restoring biological tissues without permanent scaffold retention, there are challenges in realizing longterm implant success in the absence of biomechanical stimuli in the ear site to encourage and direct tissue regeneration. [16]When this is coupled with wound contraction forces, [17] the resorption of soft tissues, chondrocyte hypertrophy and calcification may oc-cur during degradable implant resorption, leading to poor longterm aesthetic outcomes unless scaffold properties are carefully controlled. [18,19]his study presents the 3D printing of a clinical preferred implantable polymer, HDPE, with micron-scale scaffold design and fabrication capacity achieved using a technology termed melt electrowriting (MEW). [20]As an emerging 3D printing technology, this is also the first time that HDPE has been processed using MEW.This is important as the microfiber scaffold resolutions resulting from MEW are proven to be advantageous in other biomedical applications. [21,22]This has been attributed to the properties imparted by the microfibers, which are typically between 5-50 μm compared to several hundred microns produced using fused filament fabrication (FFF) 3D printing techniques. [8,23][28][29] This new MEW approach presented here implements a filament-driven printhead system, as compared to the traditional pressurized syringe approach, and a heated collector is leveraged to produce high porosity HDPE scaffolds designed to mimic the mechanical and morphological properties of auricular cartilage.

Optimization of MEW Parameters for HDPE Scaffold Fabrication
Microfiber scaffolds fabricated from high-density polyethylene (HDPE) filament were successfully fabricated using a filamentdriven MEW printhead system.Briefly, adaptations were performed to the printhead and collector of an existing open-source Voron 3D printer to provide electrical isolation between the heating and temperature monitoring components (Figure 1A).The inclusion of a MACOR block to electrically insulate the nozzle but maintain thermal conductivity ensures that in the event of an electrical discharge between the collector plate at high voltage and the grounded nozzle, the printhead and operating components would not be damaged. [20]The use of glass slides as a printing substrate further provided protection against spontaneous electrical discharge between the collector and nozzle whilst printing at low (sub-3 mm) collector distances.Maintaining low working distances (1-3 mm) was critical to optimize the deposition of the polymer whilst still in molten phase, as premature solidification is known to destabilize fabrication reproducibility. [30]he printhead described here has a cylindrical compartment "hot end" during which the filament underwent melting just prior to extrusion through the nozzle.Optimization of the extrusion feed rate, the rate at which the motors feed filament into the cylindrical chamber, was essential to ensure a continuous supply of polymer in the reservoir was available.In the event that the feed rate is set too low, an initial volume of molten polymer was deposited into fibers, shown in Figure 1B, but fiber production was quickly interrupted due to the lack of material being fed into the printhead.This unstable printing regime was overcome by increasing the extrusion feed rate, E, controlled by Melt electrowriting of HDPE scaffolds using A) a filament-driven printhead system, described elsewhere, [20] which is subject to the fill level of the heated reservoir decreasing when the filament feed rate is insufficient.B) Fibers produced after filling the reservoir but then setting the extrusion feed rate to zero, resulting in rapid jet disruption.C) Micrographs of microfibers printed with increasing extrusion feed rate and D) measured fiber diameters (reported as average ± standard deviation, n = 8).E) Images of fiber printing using various extruder and print bed temperatures.F) Micrograph of defect-free HDPE scaffolds with 60.5 ± 2.6 μm fibers, alongside PCL scaffolds printed for comparison with 28.0 ± 2.0 μm fibers.G) Representative stress-strain curves for HDPE and PCL scaffolds, normalized to cross-sectional area, and H) representative images of scaffolds during tensile testing.
the dimensionless G-code parameter, such that fibers of increasing diameter and homogeneity were obtained.Figure 1C depicts representative fibers printed using increasing extrusion feed rates that apply increasing amounts of pressure to the molten polymer within the printhead and therefore fabricate fibers with increasing diameter.These extrusion feed rates (E0-5000) are equivalent to input filament flow rates of 0.28 μL min −1 (E500), 0.56 μL min −1 (E1000), 1.12 μL min −1 (E2000) and 2.8 μL min −1 (E5000).Fibers ranging from 38.5 ± 23.9 μm (E500) up to 139.1 ± 3.0 μm (E5000) were produced, enabling the optimization of fiber diameters over a wide range by adjusting the extrusion feed rate alone (Figure 1D).Fiber pulsing [31] and other significant instabilities were observed whilst printing at low flow rates (E500), leading to high variability in measured fiber diameters.
Several other MEW parameters required adjustment to improve the consistency of the deposited microfibers.Importantly, the extruder temperature at which the filament was melted plays a significant role in ensuring the material viscosity is sufficiently low to enable extrusion of fibers in a highly reproducible fashion (Figure 1E). [29]Fiber pulsing and other instabilities were observed at temperatures below 250 °C, whilst at 250 °C, the fibers were consistent and homogeneous.The bed temperature is known to decrease the thermal gradient between the extruded molten polymer and collector plate to assist with fiber adhesion to the collector, as well as optimize the point at which the fiber solidified. [30]For the fabrication of high melting temperature polymers such as HDPE, ensuring that the polymer within the jet remains in a molten state until deposition is of critical importance for generating sufficient inter-layer bonding for functional scaffold production. [32]Therefore, the bed temperature set to just below the melting temperature of the material has been established as the best practice for ensuring and controlling fiber bonding.Here, 110 °C was determined to be the ideal bed temperature for HDPE with a melting temperature of ≈132 °C, whilst scaffolds printed at lower temperatures readily delaminated under handling (Figure S2, Supporting Information).Additional parameters were selected according to routine MEW optimization, such as the nozzle-to-collector distance that was set to 1 mm and print speed that was set to 200-or 400-mm min −1 for E1000 and E2000, respectively.Figure 1F depicts the successful fabrication of HDPE scaffolds with an average fiber diameter of 60.5 ± 2.6 μm (n = 8) and porosity of ≈96%.Scaffolds were fabricated using a crosshatch pattern with 500 μm spacing and up to five layers high (five layers of orthogonal fiber pairs).These were compared to PCL scaffolds printed using the same design and routine fabrication conditions, however, printed on a conventional air pressuredriven MEW system. [33]Medical-grade PCL scaffolds exhibited tensile mechanical performance comparable to similar scaffolds reported elsewhere in the literature [34,35] (Figure 1G) whilst the HDPE scaffolds were significantly stronger.The ultimate tensile strength of the HDPE scaffolds was 21.8 ± 1.3 MPa compared to just 3.9 ± 0.1 MPa for PCL scaffold, a 5.7-fold increase (p = 1 × 10 −5 , n = 3).The stiffness of the HDPE scaffolds was also 13 times higher than PCL scaffolds (14.7 ± 1.9 MPa compared to 1.1 ± 0.0 MPa, p = 0.0002, n = 3).The elastic modulus of the HDPE material was similarly 13 times higher than the PCL (10.4 ± 1.3 MPa compared to 0.8 ± 0.0 MPa, p = 0.002, n = 3) when calculated based on the stress-strain curve that was normalized for the difference in fiber diameter between scaffold groups (60.5 ± 2.6 μm for HDPE compared to 28.0 ± 2.0 μm for PCL).The HDPE scaffolds failed at 4.1 ± 0.8% tensile strain, whilst the PCL scaffolds exhibited significantly higher ductility than the HDPE scaffolds, characterized by their ability to undergo uniform loading up until beyond the maximum strain range tested (250%) (Figure 1H).However, the utility of such plastic deformation behavior has little utility in tissue engineering scaffold design since devices would not be used in such biomechanical environments.
The filament-driven printhead system offers unprecedented control over the flow rate of polymer being deposited using the MEW system.Using this system, the extrusion feed rate can be directly controlled by input G-code parameter, E, multiplied by a constant, k, that relates the speed of the filament motors (depicted in Figure 1A and described in full in the Supporting Information) with the input speed of filament being controlled by stepper motors feeding the material into the printhead. [20]Syringe-based MEW printhead systems typically rely on gas pressure applied to the material to facilitate extrusion. [36]However, the viscoelastic properties of the material therefore dictate the resultant extrusion flow rate achieved, subject to a myriad of compounding factors such as temperature, shear thinning behavior, and nozzle geometry. [29,37]Here, the flow rate is largely independently controlled by the filament feeding mechanism, regardless of material properties or other operational parameters due to the ability to instigate large forces on the material to maintain the programmed extrusion speed.
Due to the electrohydrodynamic and gravitational pull on the polymer through the nozzle during MEW fiber production, the material reservoir can be depleted faster than the mechanical filament feed fills it, causing instability, fiber pulsing, and discontinuity in fiber production as the nozzle reservoir runs out.In order to continuously melt and extrude the polymer in this printhead configuration, the small reservoir of molten polymer within the printhead must be continuously supplied with the filament at a proportional flow rate to the deposited fiber.The rate of filament extrusion, controlled by the G-code parameter, E, must therefore be strategically selected to both optimize the morphology of printed fibers, and ensure that supply of filament to the printhead reservoir is sufficient to enable continuous printing over several hours (Supporting Information).Calculation of the volume of input filament relative to extruded fibers at varying extrusion rates revealed that, intuitively, the reservoir fill level declines if insufficient filament feed rate is selected (Figure 2A).Where a significant mismatch in input flowrate compared to output flowrate of printed fibers was observed, for example using E5000 to produce 140 μm fibers, the reservoir fill level dropped to 0% after just 15 min.Therefore, the maximum print time before the reservoir fill level drops to 50%, for example, is constrained by any disparity between input and output flow rates (Figure 2B) but can be readily mitigated by better predicting the ideal extrusion rate for each target fiber diameter.
The ideal extrusion rate is calculated directly via Equation 2. Figure 2B includes asymptotic limits for each of the printing regimes where the print time is infinite for each ideally calculated extrusion rate value, E ∞ .The calculations to generate these graphs require iteration since they rely on experimental measurement of fiber diameters that is proportionally affected by the extrusion rate, as well as other parameters not captured in this model.The graphs therefore provide a useful estimate of the ideal extrusion rate value for printing fibers with a diameter tolerance of ±10 μm within a printing window of 5 h.The precision optimization of E ∞ is required when considering industrial scalability of filament-driven MEW technology and the ability to print, uninterrupted, for considerably longer timescales than achievable with a finite supply of polymer using syringe-based systems.

Thermal Degradation of HDPE During Extrusion
One substantial improvement to MEW fabrication offered by the filament printhead is the significant reduction in heat exposure experienced by printed materials compared to syringe-based systems.In these latter configurations, the entire volume of printing material, often up to 3 mL, is retained in molten state at high temperatures between 30 and 50 °C above the material's melting point over the entire duration of multi-hour and even multi-day prints, subjecting the material to substantial thermal energy, inducing degradation. [27,34,38]Conversely, the filament-driven printhead used in this present study has a heated polymer residence time within the reservoir of an estimated 185 min for an extrusion rate of E500 (0.28 μL min −1 ), and as low as 18.5 min for E5000 (2.8 μL min −1 ).
To examine the influence of heat exposure on HDPE, degradation analysis was performed by heating samples to 200 and 250 °C for 6 h, representative of a typical MEW print.estimated residence time of polymer traveling through the filament printhead at E2000, and 6 h of continuous heating that would be experienced by polymer in a syringe printhead in a traditional MEW device.After 6 h heating at 250 °C, a mass loss of 8.7% was reported, compared to just 3.7% after just 45 min heating (Table 1).Whilst degradation was largely mitigated by heating to just 200 °C for 45 min, inconsistent fiber deposition was observed during printing (Figure 1E), highlighting that balancing optimized viscoelastic properties for consistent fiber production with thermal degradation mitigation must be considered.Further evidence of the impact of heating on the polymer structure is demonstrated in the DSC results (Figure 3B), whereby the crystallinity of the HDPE shifted from 45.7% to 52.1% after 6 h heat exposure, which was not observed in samples that had not exposed to this isothermal step.Visually, the polymer appears yellow/brown after 6 h heat exposure, characteristic of degradation, [39] compared to the cool white shade of the nonheated polymer (Figure 3C).These quantitative and qualitative findings corroborate with previous reports of increased stiffness, brittleness, and loss of ductility of materials as a result of extended heat exposure during MEW. [24,27,34,38]The ability to preserve the thermal properties of materials due to significantly less heat exposure presents a significant opportunity to process a wider range of polymers at high temperatures than previously achievable.

Biomimetic Auricular Scaffold Designs
Auricular scaffolds were designed using the gold standard MED-POR auricular surgical implant as a template, intended to mimic the auricular cartilage structure within an adult ear. [40]EW scaffolds were successfully fabricated based on this design (Figure 4A; Video S1, Supporting Information) using 500 μm fiber spacing and five crosshatch layer pairs.Using the MED-POR design as the template, scaffolds matched to the anatomy of two photographs of adult ears were also designed and fabricated using the same processing parameters (Figure 4B,C), alongside an additional standard model ear but using three infill patterns, 200, 500, and 1000 μm crosshatch fiber spacing (Figure 4D) corresponding to the high stiffness cartilage, low stiffness cartilage and connective tissue between cartilage structures, respectively, per published literature on native auricular cartilage mechanical properties. [41]Connecting diagonal lines are present throughout  the scaffolds, most notably visible in Figure 4D(ii-iv) opposing the linearity of the vertical and horizontal fibers comprising the crosshatch patterns.These are generated by the collide-and-turnbased algorithm for generating the print path for complex and interconnecting regions of the auricular design and are inherent to the MEW process that uses one continuous fiber to print the entire structure. [42]his semi-automated scaffold design and manufacturing methodology created scaffolds that conformed to the unique anatomical features of the example ears.For example, the scaffolds produced were able to accurately represent the substantial variation in curvature of the helix between the two ears, the Cshaped outer rim of the auricle (Figure 4B,C) and offers a promising and rapid avenue for augmenting generic surgical implant designs toward patient-matched product manufacturing.
This research has established a proof-of-concept methodology for manufacturing biomimetic auricular surgical implants to fabricate, for the first time, melt electrowritten HDPE microfibers.Several opportunities exist for further development and refinement toward optimizing these scaffolds as surgical implant products.One aspect is adjusting the scaffold thickness relative to the mechanical properties to both withstand wound contraction forces without compromising on the natural "look and feel" of the implant, a significant consideration for patient satisfaction and overall aesthetic outcome. [43]Further studies are required to closely investigate the relationship between scaffold stiffness, contraction, and resulting shape fidelity.Additionally, future iterations of the implants may benefit from the inclusion of projection blocks to ensure firm and stable projection of the ear from the side of the head, improving implant longevity and reducing the risk of dislocation or deformation over time. [44]The biocompatibility of microfiber HDPE provides an avenue for future enhancements in tissue integration and vascularization.Medicalgrade HDPE filament, with low impurities, is required to meet regulatory standards, and while there is a limited number of com-mercial HDPE filaments available due to the inherent challenges of processing them using FFF, this study contributes to the growing evidence base for fabricating medical products from HDPE and it is anticipated that this will drive the development and increase the availability of medical-grade filament feedstock supplies in the future.Finally, surface modifications of the HDPE material may be explored to enhance the ingrowth of native tissues, fostering a more robust and natural bond between the implant and the host tissue. [45]These adaptations may also improve vascularization within the implant, promoting better nutrient and oxygen supply to the integrated tissues, and enhancing their survival and function.The encouraging initial findings reported in this study, coupled with the opportunity for future progression toward clinical translation, suggest that MEW-based HDPE implants present an exciting new avenue for auricular reconstruction.

Conclusion
This study reports the 3D printing of a clinically proven polymer for auricular surgical implants that addresses ongoing challenges associated with permanent surgical implants.The key innovation of this study lies in the successful utilization of a filamentdriven printhead MEWron 3D printer to melt electrowrite HDPE for the first time, a task not previously achievable with conventional syringe-based MEW systems due to pneumatic pressure limitations for high-viscosity materials as well as thermal degradation.Here, we demonstrated that the rapid application of heating and subsequent cooling of small volumes of the printing material using the filament-driven system reduced the thermal degradation of HDPE to just 3.7% mass loss compared to 8.7% mass loss when heated for 6 h per routine practice for syringe-based MEW systems.The HDPE scaffolds fabricated using the filament-driven system exhibited a remarkable 5.7-fold increase in ultimate tensile strength and a 13-fold increase in stiffness compared to conventional MEW PCL scaffolds, offering a significant opportunity to produce higher strength and nondegradable scaffolds using this technology.These mechanical enhancements may greatly improve the ability to translate scaffold products for higher-strength tissue engineering and surgical applications, with improved durability and functionality in various biological contexts.This work paves the way for an improved approach to designing and fabricating patient-specific implants, potentially revolutionizing the outcomes for auricular cartilage reconstruction procedures.The filament-driven printhead MEW system thus holds a promising future for further enhancements in the field of biomimetic implants and beyond.
Image Acquisition and Scaffold Design: Photographs of adult human ears were captured with permission using a smartphone camera (Samsung S20).A spine outline of the MEDPOR implant as a template [40] was morphed to conform to the anatomy in each image in Photoshop (Adobe Inc., USA).Each morphed spline outline was filled with black and exported to a 512 × 512 px binary bitmap (BMP) image.The image was then imported into a bespoke G-code generation program reported previously for interpreting images into MEW-specific G-code following a continuous fiber path with user-defined orientation and fiber spacing. [42]Scaffolds with 500 μm spacing were programmed for each auricle design, as well as a multi-porosity scaffold using regions of 200, 500, and 1000 μm spacing inspired by varying stiffness regions of native auricular cartilage. [41]ectangular scaffolds 10 × 25 mm in dimension were also fabricated with crosshatch architecture and 500 μm spacing between parallel fibers.All scaffolds were printed with five pairs of orthogonal crosshatch layers, totaling ten individual layers.
Melt Electrowriting (MEW): Scaffold fabrication was performed on two MEW devices: a MEWron system fitted with a filament-driven printhead, described previously, [20] and a custom-built system with traditional air pressure and syringe-based printhead, also described previously. [33]For 3D printing using the MEWron, a HDPE filament was fed into the printhead set at 200-250 °C, fitted with a 0.15 mm brass nozzle (E3D Online, UK).The collector plate temperature was set to room temperature or 110 °C with the distance set as 1 mm between the nozzle and a glass slide (76.2 × 50.8 × 1 mm), placed on the collector onto which scaffolds were printed.The high voltage applied to the collector plate was set to 3.5 kV.Several parameters were controlled from the G-code, including the extrusion rate, E, altered between E0-E5000, the print speed altered between 100-600 mm min −1 , and scaffold design paths, designed using custom Gcode generation described above or simple parallel lines, spaced 0.5 mm apart.To ensure the printhead reservoir was full at the beginning of each print, the filament was manually pushed into the printhead until a thick fiber was produced, followed by the fabrication of two stabilization lines at E10000 before commencing the scaffold print path at a lower extrusion rate.PCL scaffolds using the same G-code were printed on a traditional syringe-based system as controls using 1 bar air pressure, 3 mm collector distance, 600 mm min −1 print speed, and 4.5 kV high voltage applied to the nozzle.
Extrusion Prediction: A series of equations were developed to predict the rate of polymer extrusion through the filament-driven printhead, based on the input extrusion rate G-code factor, E, extruder gear ration constant, k, filament diameter, ϕ filament , reservoir volume, V res , programmed print speed, v, and measured fiber diameter, ϕ fiber .A full derivation is provided in the Supplementary Information.
The reservoir starts full (100% fill level) by manually pressing the filament into the printhead.After t mins, the fill level of the reservoir is reduced at the rate of fiber extrusion onto the collector and increased at the rate of material being fed in via the filament extruder motors, which is expressed in Equation 1.
V res (1)   Similarly, the ideal value for the extrusion rate, E ∞ , to ensure continuous filament supply matching the flow rate of fiber production can be calculated by simply equating input and output flow rates, described in Equation 2. 2 Mechanical Testing and Imaging: Rectangular scaffolds fabricated from either HDPE or PCL underwent tensile mechanical testing in triplicate (Tytron 250, MTS) using a 250 N load cell under displacement control.Scaffolds were loaded at a rate of 2 mm s −1 , corresponding to a strain rate of 10% strain/sec.Stress-strain curves were calculated by translating displacement to strain and force to stress by normalizing to the crosssectional area of the scaffold fibers, measured using stereomicroscopy (VHX-7000, Keyence, USA) in replicate (n = 8).The tile scanning microscope feature was used to visualize entire auricular scaffolds in a single image.Porosity was estimated using the fiber diameter to calculate the volume of fibers divided by the total volume of the scaffold.
Thermal Analysis and Degradation: The thermal properties of the HDPE and PCL were measured using digital scanning calorimetry (DSC) (NETZSCH 204 F1 Phoenix) and thermogravimetric analysis (TGA) was performed on a simultaneous thermal analyzer (NETZSCH STA 449F3 Jupiter).Heating and cooling cycles were performed between 25-250°C, with and without a 6 h isothermal step at 250 °C.The melting temperature and crystallinity were measured from the first cycle to characterize initial melt behavior and the second cycle following the isothermal step to characterize crystallinity after long-term heating.Isothermal analyses at 200 and 250 °C were performed for 6 h before heating to 600 °C at a rate of 10 °C min −1 to analyze degradation behavior during long-term exposure to heating within syringe-based MEW printheads.
Patient Consent Statement: Informed consent was received for the publication of images of distinctive body features.

Figure 1 .
Figure 1.Melt electrowriting of HDPE scaffolds using A) a filament-driven printhead system, described elsewhere,[20] which is subject to the fill level of the heated reservoir decreasing when the filament feed rate is insufficient.B) Fibers produced after filling the reservoir but then setting the extrusion feed rate to zero, resulting in rapid jet disruption.C) Micrographs of microfibers printed with increasing extrusion feed rate and D) measured fiber diameters (reported as average ± standard deviation, n = 8).E) Images of fiber printing using various extruder and print bed temperatures.F) Micrograph of defect-free HDPE scaffolds with 60.5 ± 2.6 μm fibers, alongside PCL scaffolds printed for comparison with 28.0 ± 2.0 μm fibers.G) Representative stress-strain curves for HDPE and PCL scaffolds, normalized to cross-sectional area, and H) representative images of scaffolds during tensile testing.

Figure 2 .
Figure2.A) Estimation of the fill level of the printhead reservoir whilst printing fibers using various programmed extrusion rates.B) Estimation of the maximum print time until the reservoir is empty based on increasing extrusion rates.

Figure 3 .
Figure 3. A) Thermogravimetric analysis (TGA) of HDPE samples, heated to 200 or 250 °C and held for 6 h (dashed lines) before heating to 600 °C.Thermal degradation reported as mass loss during temperature ramp as well as (inset) isothermal measurement segments.B) Digital scanning calorimetry (DSC) measurement of the melting peak of HDPE and calculation of crystallinity during initial melting to 250 °C (green) as well as after 6 h isothermal at 250 °C (red).C) Composite image of HDPE in a syringe (traditional MEW) before (bottom) and after 6 h heating at 250 °C (top).

Figure 4 .
Figure 4. Melt electrowritten HDPE scaffolds resembling auricular surgical implants.A) Standard design using the MEDPOR implant as the design template that was then augmented to match the morphology of two adult ears (B(i) & C (i)), depicted alongside images of the ears (B(ii) & C(ii)).Microscopy images of a multi-stiffness scaffold (D(i)) with three regions of varying infill pattern (D(ii-iv)).

Table 1 .
Comparison of mass loss, as measured using a simultaneous thermal analyzer (STA) during isothermal heating of HDPE for 6 h.

Table 1
in combination with Figure 3A, reports the measurement of mass loss of HDPE after 45 min heat exposure, corresponding to the