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Design and fabrication of a hybrid alginate hydrogel/poly(ε‐caprolactone) mold for auricular cartilage reconstruction

Abstract The aim of this study was to design and manufacture an easily assembled cartilage implant model for auricular reconstruction. First, the printing accuracy and mechanical properties of 3D‐printed poly‐ε‐caprolactone (PCL) scaffolds with varying porosities were determined to assess overall material properties. Next, the applicability of alginate as cell carrier for the cartilage implant model was determined. Using the optimal outcomes of both experiments (in terms of (bio)mechanical properties, cell survival, neocartilage formation, and printing accuracy), a hybrid auricular implant model was developed. PCL scaffolds with 600 μm distances between strands exhibited the best mechanical properties and most optimal printing quality for further exploration. In alginate, chondrocytes displayed high cell survival (~83% after 21 days) and produced cartilage‐like matrix in vitro. Alginate beads cultured in proliferation medium exhibited slightly higher compressive moduli (6 kPa) compared to beads cultured in chondrogenic medium (3.5 kPa, p > .05). The final auricular mold could be printed with 300 μm pores and high fidelity, and the injected chondrocytes survived the culture period of 21 days. The presented hybrid auricular mold appears to be an adequate model for cartilage tissue engineering and may provide a novel approach to auricular cartilage regeneration for facial reconstruction. © 2018 Wiley Periodicals, Inc. J Biomed Mater Res Part B: Appl Biomater 107B: 1711–1721, 2019.


INTRODUCTION
Auricular reconstruction following trauma (e.g., burns), cancer, or congenital anomalies is a very challenging procedure. To date, several treatment options exist to reconstruct a deformed auricle including synthetic implants (e.g., Medpor) 1-3 and autologous reconstruction using rib cartilage. [4][5][6] However, both techniques have several limitations including donor site morbidity, 7,8 risk of implant exposure, 1,9 and surgical complexity. 10,11 Tissue engineering, in combination with novel biofabrication strategies, is a promising solution to engineer auricular implants with patient-derived donor cells. [12][13][14][15] These biofabricated auricular constructs could ultimately function as patient-specific implants for the reconstruction of a deformed auricle.
Successful engineering of a clinically relevant auricular cartilage implant requires a scaffold that provides a threedimensional (3D) environment to support cell and tissue Additional Supporting Information may be found in the online version of this article. growth. The optimal scaffold ideally matches the native properties of cartilage extracellular matrix (ECM) and is mechanically strong, porous, and biodegradable. 16,17 Unfortunately, creating such a construct is still a major challenge in the field of tissue engineering. In addition, efforts to translate engineered cartilage to the clinic have been hampered by problems associated with scaling up. 18 3D printing can aid in both the production of 3D scaffolds that closely mimic the ECM of cartilage in terms of biomechanical properties and in producing larger, anatomically shaped tissues such as whole auricles. One strategy to engineer cartilage using 3D printing is to directly deposit the cells in a layer-by-layer fashion. The "bioink" that is required to directly deposit such cells consists of a synthetic, or more commonly, natural hydrogel mixed with cells. 19 Although hydrogels can closely resemble the native ECM of cartilage, the mechanical strength of such gels is low. 20 Therefore, hydrogels often require a synthetic supporting scaffold or microfibers that temporarily increase the mechanical properties. 21,22 These supporting scaffolds can be co-printed with the hydrogel to form an integrated scaffold construct.
Another common strategy to engineer cartilage using 3D printing is to print the supporting scaffold and manually load the cell-hydrogel mixture onto the 3D-printed scaffold, traditionally known as "top-down" approach. 23 Using this strategy, cells must deposit their own ECM on the prefabricated scaffold. A commonly used polymer for creating such a supporting scaffold is poly-ε-caprolactone (PCL). PCL is a promising biopolymer for tissue engineering due to its tuning possibility, chemical versatility, and strength, 24 and has been used in a clinical setting.
In a previous study, our group introduced the cage construct. 25 In that study, we developed a 3D-printed construct with a central core consisting of a natural polymer with cells, and an outer synthetic PCL cage for mechanical support. In the current study, we aimed to take this construct one step further by creating a clinically relevant and simplified twopart scaffold construct that can be easily assembled in the shape of human auricular cartilage. First, we tested the mechanical properties of both 3D printed PCL scaffolds and alginate hydrogel beads and discs and tested the survival of isolated chondrocytes inside the gel alone, with and without additional chondrogenic stimulation. Subsequently, we combined the two concepts by designing and 3D printing a twopart PCL mold in the shape of human auricular cartilage and injected this construct with a mixture of alginate and chondrocytes before culturing the final construct in vitro.

Cell isolation and culture
Cell expansion. Cell isolation and culture of chondrocytes was performed as described previously. 25 Briefly, cartilage was obtained from the ears (n = 4) of sacrificed Dutch milk goats acquired from a local abattoir. The dissected cartilage was treated with 0.2 wt % type II collagenase (Roche diagnostics, Almere, the Netherlands) under gentle agitation at 37 C for 16 h to isolate the chondrocytes from the cartilage tissue. The digested tissue was subsequently filtered using a 40 μm-mesh to remove residual ECM, pelleted using centrifugation, and washed with PBS. The resulting cell suspension was cultured in GIBCO Dulbecco's modified Eagle's medium (DMEM; Life Technologies, Thermo Fisher Scientific, Waltham, MA, USA) supplemented with 10% fetal bovine serum (FBS; Hyclone-Thermo Scientific, Thermo Fisher Scientific, Waltham, MA, USA), 1% PSF (10,000 U/mL Penicillin, 10 mg/mL Streptomycin, 25 mg/mL Amphotericine B (Fungizone; Millipore Sigma, St. Louis, MO, USA), and 50 μg/ mL L-(+) ascorbic acid (Merck, Darmstadt, Germany) (proliferation medium). When cell cultures reached 80% confluence, chondrocytes were trypsinized using 0.05% trypsin-EDTA, resuspended in cryoprotective medium (Recovery Freezing Medium, Life Technologies, Thermo Fisher Scientific, Waltham, MA, USA), and stored in liquid nitrogen until further use. All chondrocytes taken from liquid nitrogen were cultured up to passage three (P3) in proliferation medium to allow sufficient cells for three-dimensional (3D) culture in alginate hydrogel.
3D culture in alginate hydrogel beads. Sterile alginate solution was prepared by dissolving ultrapure sterile sodium alginate (MW 150-250 kDa, Pronova SLG100, Novamatrix, FMC Biopolymer, Sandvika, Norway) in 0.9 wt % Sodium Chloride (NaCl). The resulting alginate solution was vortexed at high speed and subsequently kept at room temperature for at least 1 h to minimize bubbles in the solution. For 3D alginate culture, cultured auricular chondrocytes were slowly suspended at a density of 4 × 10 6 cells/mL in 3 wt % alginate solution and transferred into 10 mL sterile syringe. The suspension was slowly passed through a 21-gauge (21G) needle to produce small drops, which fell into a sterile 102 mM CaCl 2 solution to allow instantaneous gelation. Following gelation, the beads were allowed to further gelate for 10 min in CaCl 2 solution before being washed with 0.9 wt % NaCl and transferred to a 24-well plate.
Auricular cartilage mold. An auricular cartilage mold, based on the parametric model described in the study by Bos et al., 26 was designed in CATIA V5 software (Dassault systemes, Vélizy-Villacoublay, France). The two-part mold was processed in the custom software (see above) to obtain a controllable G-code for 3D printing. Using the optimal porosity from the above experiment, the auricular mold was then printed in two parts on the 3DDiscovery printer using the same settings as described previously. Following 3D printing, the molds were disinfected in 70% ethanol for 1 h, washed 3 times with sterile PBS, and air-dried in a sterile incubator. Following disinfection, all mold parts were transferred to individual petri dishes and coated with sterile gelatin (10%, Merck KGaA, Darmstadt, Germany) to block the pores and generate a sealed mold. After gelation of the gelatin solution at room temperature, 400 μL of the cell-alginate mixture (same suspension as described in 3D culture in alginate hydrogel beads) was injected into the bottom part of the mold, crosslinked for 2 min using 102 mM CaCl 2 , and directly closed by placing the top part of the mold. Upon subsequent immersion into proliferation medium at 37 C, the gelatin liquefied and vacated the pores to allow nutrient and oxygen exchange through the mold pores. Figure 1A provides a complete overview of the fabrication process. The complete auricular implant models (Fig. 1B) were cultured in vitro in proliferation medium at 37 C, 5% CO 2 for 21 days and processed for LIVE/DEAD staining.

Microscopy
Optical microscopy (5× Objective, Nikon Eclipse LV100ND, Tokyo, Japan) was performed to assess the microscopic architecture of the 3D printed PCL scaffolds. Scaffolds were imaged from the top, bottom, and side view and processed in accompanying software (NIS-Elements, Nikon). The overall structure of the 3D printed PCL scaffolds was assessed with a stereoscopic microscope (Nikon SMZ-10) with a digital camera (Nikon DXM1200F) and fiber optic light source (Schott, KL1500-T).
A Light sheet microscope (UltraMicroscope II, LaVision BioTec GmbH, Bielefeld, Germany) equipped with a white light laser (SuperK EXTREME, NKT photonics; excitation 400-800 nm) and cMOS camera (Neo, Andor) and a Nikon Eclipse 80i confocal microscope were used to image chondrocyte-laden alginate beads and chondrocyte-laden alginate taken out of the auricular implant model respectively. The beads were placed in a holder and submerged in PBS during imaging. In general, the light sheet microscope has good optical sectioning capabilities allowing Z-plane imaging of intact hydrogels. However, light only penetrates a few hundred micrometers into transparent tissues. 27 Cell viability analysis and quantification A LIVE/DEAD viability/cytotoxicity kit (L3224; Thermo Fisher Scientific) was used to assess cell viability in chondrocyte-laded alginate beads according to manufacturer's instructions. Briefly, chondrocyte-laden alginate beads were washed in PBS, followed by incubation in 0.5 μL/mL Cal-AM and 2 μL/mL Eth-D (Thermo Fisher) in PBS at 37 C for 1.5 h. Beads were washed again in PBS and imaged using light sheet or confocal microscopy. Live and dead cell quantification was performed in ImageJ v1.47 for Mac using the "process > find maxima" function.
Cell quantification in alginate beads cultured in proliferation and chondrogenic medium was performed using DAPI staining. Unfixed chondrocyte-laden alginate beads were covered in Tissue-Tek, snap frozen, and cut in 20 μm sections using a cryostat. Following the application of mounting medium (Fluoroshield Mounting Medium with DAPI; Abcam), beads (n = 3 per group) were imaged on a Nikon Eclipse 80i confocal microscope. Cells were counted in NIS-Elements AR software (v 3.2; Nikon Instruments Europe B.V). Both the total number of cells and percent (%) cells per bead were calculated per section.

Histology and immunohistochemistry
For histology and immunohistochemistry (IHC), unfixed chondrocyte-laden alginate beads were covered in Tissue-Tek (Thermo Fisher), snap frozen in liquid nitrogen, and cut in 14-20 μm sections at −20 C using a cryostat. Sections were mounted on Superfrost Plus Gold microscopic slides (Thermo Scientific) and stained with Alcian Blue (pH 1.0) for histological examination or processed for IHC. For IHC, sections were incubated for 2 h with monoclonal antibodies against type II collagen (II-II6B3, Developmental Studies Hybridoma Bank) or type VI collagen (MAB3303, EMD Millipore). Following incubation, sections were washed in PBS and incubated for 30 min with EnVision (Dako). Finally, sections were incubated with AEC substrate for 10 min and visualized using microscopy. Postprocessing was undertaken using NIS Elements software (Nikon Instruments Europe B.V.).
Control cartilage samples derived from goat ears were embedded in paraffin using standard histological techniques, cut in 5 μm sections, and stained using Alcian Blue, Mayer's Hematoxylin-Eosin, type II and type VI collagen to assess glycosaminoglycan distribution, morphology, and FIGURE 2. CAD view, gross view, and microscopic views of 3D-printed PCL scaffolds with varying distances between strands. S represents the sample, with the number representing the distances between strands in micrometers (μm). collagen distribution, respectively (Supporting Information Fig. S1).
(Bio)mechanical analysis PCL scaffolds. The mechanical properties of PCL scaffolds with varying distances between strands (400-1200 μm; n = 5 per group) were analyzed using mechanical compression testing according to the ASTM standard D695. The 3Dprinted PCL blocks were cut into six cuboids (5 mm wide × 5 mm wide × 10 mm high) using a new blade for each sample. An Instron 5969 machine with a 5 kN load cell was used to compress unconstrained samples between two steel plates at a rate of 1 mm min −1 to 33% strain. Five samples were tested for each scaffold design (porosity). Compressive moduli were calculated for all samples using a linear-elastic compression phase as the applied force increased from 10 to 50 N.
Alginate hydrogel. For mechanical characterization of empty and chondrocyte-laden alginate beads after in vitro cell culture, compression was performed using a Physica MCR 501 rheometer (Anton Paar, GmbH, Austria). The gap between the plates was lowered from 2 mm to 0.1 mm at a rate of 0.001 mm s −1 and the resulting normal force from the sample on the top plate was recorded. To determine the storage (G') and loss (G") shear moduli, small-amplitude oscillatory shear (SAOS) rheometry at a frequency of 0.5 Hz and with a strain amplitude of 0.5% was performed on 2 mm (height) × 20 mm (diameter) chondrocyte-laden alginate disks that were created using a 3D printed mold. All samples were tested at 37 C using PP20 stainless steel geometry (parallel plates, 20 mm diameter).
Statistical analysis. Data were analyzed using GraphPad Prism version 6.02 (Graphpad Software). Data are represented as Mean AE Standard Deviation (SD) unless otherwise indicated. Statistical analysis was performed using unpaired t-tests or one-way analysis of variance (ANOVA) with multiple comparisons. Chondrocytes proliferation and storage/ loss modulus of chondrocyte-laden alginate beads and disks were analyzed using unpaired t-tests, while ANOVA was used for comparing the compressive moduli of 3D printed PCL scaffolds. A p value of <0.05 was considered statistically significant.

RESULTS
Structural properties of 3D-printed PCL. The structural properties of 3D-printed PCL scaffolds were determined by examining the surface porosity and mechanical properties. Macroscopic analysis of the PCL scaffolds showed good printing quality (Fig. 2, gross view). However, microscopic analysis of individual PCL strands showed some variance in strand diameter over a short distance (Fig. 2, microscopic  view). In addition, the lateral view of the scaffold showed a large variety in pore width (Fig. 2, side view). Overall, the smaller the pore width, the more accurate the 3D-printed scaffold. Specifically, the S400 and S600 showed qualitatively more repeatable pores than S800-S1200 (Fig. 2).
The compressive modulus was highest in the scaffolds with the smallest pore fraction (Fig. 3A). The compressive modulus of S600 and S800 did not significantly differ from S400 (p = 0.91 and p = 0.07, respectively). To ensure pores were not blocked, the S400 and S600 scaffolds were printed with lower extrusion rates (decreasing the overall strand thickness) of 10 and 15 revs/m, respectively, as opposed to 18 revs/m for other samples. This may have caused a slight decrease in compressive modulus, as well as increase in surface pore fraction (Fig. 3C). Nevertheless, there was an overall trend that the higher the pore size and pore fraction, the lower the mechanical properties. The surface pore fraction however, calculated as the percentage of open space versus total area (Fig. 3D), was notably lower in scaffolds with smaller pores (Fig. 3C): the surface pore fraction of S400 and S600 were approximately 30%, whereas the surface pore fraction of S1000 and S1200 reached approximately 50% (Fig. 3C).

Cell viability and proliferation in alginate hydrogel beads.
Chondrocyte-laden alginate hydrogel beads were evaluated for cell survival and proliferation after 21 days of cell culture. In proliferation medium, cells displayed a rounded morphology and were dispersed throughout the alginate hydrogel beads, with more cells found at the edges of the beads (Figs. 4A and 5A). In chondrogenic medium, cells were also dispersed throughout the alginate, but in multiple cell clusters (Figs. 4A and Fig. 5B). Cell survival was high in both proliferation and chondrogenic medium (82.67% AE 2% and 82.40% AE 2%, respectively, p = 0.88, Fig. 4B). Both the total number of cells (Fig. 5C), and area of cells (%) were higher in beads cultured in chondrogenic medium compared to proliferation medium (p < 0.05).
Hydrogel biomechanical properties. The biomechanical properties of chondrocyte-laden alginate hydrogels (beads and discs) were tested using compression tests and SAOS rheometry. The overall normal force of the chondrocyte-laden alginate beads following compression is shown in Fig. 6A. Beads cultured in proliferation medium exerted higher forces (0.24 N) at 1 mm gap (~50% strain) than control beads (without cells) (0.06 N) and beads cultured in chondrogenic medium (0.07 N) following a culture period of 21 days.
We next tested the elastic shear modulus (G') of chondrocyte-laden alginate disks after 28 days of culture in different culture media. The elastic modulus of discs cultured in proliferation medium (6100 AE 1925 Pa) was higher than the elastic modulus of discs cultured in chondrogenic medium (3423 AE 1092 Pa), although the difference was FIGURE 4. (A) Representative LIVE/DEAD stains of alginate beads with chondrocytes after 21 days of culture in "proliferation" medium and "chondrogenic" medium (additional TGF-β and ITS premix). Green represents live cells, magenta represents dead cells. Blurring in the center of the "chondrogenic" alginate bead is due to limited light penetration in the 3D construct. (B) Quantification of live and dead cells in alginate beads cultured in "proliferation" and "chondrogenic" medium. not statistically significant (p = 0.104), Figure 6B. Alginate discs without cells cultured in proliferation medium showed an average elastic modulus of 4163 AE 1111 Pa (Fig. 6B).
Neocartilage formation. The production of glycosaminoglycans, type II collagen, and type VI collagen were assessed in chondrocyte-laden alginate beads (Fig. 7). In proliferation medium, chondrocytes deposited some pericellular and territorial glycosaminoglycans, while in chondrogenic medium, there was pericellular, territorial, and interterritorial deposition of glycosaminoglycans. Native auricular cartilage showed mostly pericellular and territorial glycosaminoglycans, with some interterritorial glycosaminoglycans (Supporting Information Fig. S1).  After 21 days, type II collagen produced by chondrocytes in beads cultured in proliferation medium was deposited mostly in the pericellular region. Chondrogenic medium stimulated some additional deposition in territorial and interterritorial regions (Fig. 7). Type VI collagen deposition was mostly pericellular in both proliferation and chondrogenic medium. However, large cell clusters in chondrogenic medium seemed to have formed some territorial matrix. In native auricular cartilage, type VI collagen, although most dense in the pericellular/territorial region, appeared rather dispersed throughout the whole tissue (Supporting Information Fig. S1).
Auricular implant model. To assess the feasibility of developing an auricular implant model, both PCL and alginate were combined in one construct. From the above   (Figs. 2 and 3), a porosity of 300 μm (S600) was chosen for the auricular implant model because it demonstrated both a high compressive modulus and repeatable pore geometry. The two-part auricular mold (Figs. 1 and 8A, B) could be printed with high fidelity. The final size of the assembled PCL mold was 40 mm (height) × 25 mm (width) × 6 mm (thickness). After the in vitro culture period in proliferation media for 21 days, the top mold could be removed easily without alginate sticking to the PCL. After removal (Fig. 8A), the alginate was seen to be evenly distributed in the grooves of the PCL mold. The size of the mold did not change during in vitro culture. Cell survival was high (>85%) in the alginate gel after 21 days of culture at all positions in the construct (Fig. 8C). The cells preferred the alginate as microenvironment as opposed to the surrounding PCL; very few cells attached to the PCL scaffold. Furthermore, cells were distributed evenly both throughout the model and in the alginate hydrogel.

DISCUSSION
Advances in the field of tissue engineering and biofabrication provide a possibility to overcome current issues in auricular cartilage reconstruction. One of these issues is creating an implant model that both mimics the complexity of the auricle and maintains this complexity during cellular growth and tissue remodeling. In the current study, we developed and tested a biofabricated auricular implant model in vitro, that can be used for ear reconstruction.
Appropriate selection of the biomaterials is vital for the long-term success of implants. 28 In this study, PCL was used because of its inherent strength for potential subcutaneous in vivo implantation, 13 high elasticity, 29,30 and low-melting temperature of 60 C, making it an excellent biomaterial for 3D printing. 31 3D-printed PCL scaffolds exhibited very highcompressive moduli (100-200 MPa), depending on the porosity (Fig. 3A). 32 These compressive moduli are approximately 100 times higher than that of native auricular cartilage, 33,34 and are similar to the clinical synthetic implant material porous polyethylene (Medpor), 10,34 which is widely used for auricular reconstruction. 3,35 The high stiffness of the PCL is both advantageous for protection of the hydrogel as well as shape preservation following possible implantation. 36 Scaffold porosity is an important aspect in biofabrication-based tissue engineering. 37 For example, Medpor constructs, with pores greater than 100 μm, allows soft tissue and vascular ingrowth when implanted in patients. 38 We found the optimal porosity in terms of mechanical properties, 3D design, and PCL material, to be 300 μm (S600) (Fig. 3C). Although, 3D-printed scaffolds with smaller pore sizes (S400) seemed more accurate (Fig. 2), some pores were blocked, which would be expected to block nutrient diffusion in the hydrogel. In addition, smaller pore sizes have been shown to promote less tissue ingrowth than larger pores. 39 With regards to nutrient diffusion in the alginate hydrogel, beads cultured in both proliferation medium and chondrogenic medium displayed high cell survival after 21 days in vitro (~83%). There are various reasons why cell survival is not 100%, including chemical crosslinking with calcium chloride, 40 alginate viscosity, 41 and cell culture conditions. 42 Stimulation with chondrogenic medium did not increase the percentage of cell survival but did significantly increase cell proliferation and cluster formation (Fig. 5). Cell cluster formation in 3D hydrogels has been shown previously, [43][44][45] but the mechanism by which cell clusters are formed is poorly understood. 46 Cell clustering could be induced by stimulation with TGF-beta present in our chondrogenic medium. However, this phenomenon is mostly seen in osteoarthritic models. 47,48 Nevertheless, a study by Cavo and colleagues 49 showed that MCF-7 cells show high-cell proliferation, associated with cluster formation, in soft alginate hydrogels. In addition, Mhanna and colleagues 50 explain that alginate inhibits cell migration and forces cells to proliferate in nodes resulting in cell clusters. In this study, the compressive modulus of alginate beads stimulated with chondrogenic media was lower while the proliferation rate was higher. Therefore, proliferation rate might not only be related to hydrogel stiffness, but also cluster formation.
The stiffness of in vitro cultured chondrocyte-laden alginate beads was low (~5 kPa, Fig. 6B) compared to native auricular cartilage (~1 MPa 34 ). Although not statistically significant, there was a noticeable difference in the elastic modulus of beads cultured in proliferation medium (~6 kPa) compared to chondrogenic medium (~3.5 kPa). Although, we expected that stimulation with TGF-beta would increase the mechanical properties of the beads by increasing chondrocyte proliferation and matrix synthesis, the opposite was found. This negative effect of continuous TGF-beta stimulation has also been shown by others, 51-53 which could suggest that transient exposure of TGF-beta is important to improve tissue formation. 52 A study by Bhujbal and colleagues 42 identified several factors that may influence the overall mechanical stability of alginate beads, including storage solution and cell load. They showed that the mechanical properties of beads cultured in DMEM decreased significantly. This might also explain the low elastic moduli of our chondrocyte-laden alginate beads in vitro. In addition, higher cell load could explain a lower elastic modulus in chondrocyte-laden alginate beads stimulated with TGF-beta (Fig. 5).
ECM deposition was more prominent in chondrocyteladen alginate beads cultured in chondrogenic medium. Both glycosaminoglycans and type II collagen deposition were higher in beads cultured in chondrogenic medium, which is consistent with previous data. [54][55][56] To determine whether the pericellular matrix of auricular chondrocytes was affected by TGF-beta stimulation, we performed type VI collagen immunostaining. Type VI collagen was weakly positive in the pericellular regions of both chondrocytes stimulated with proliferation and chondrogenic medium, which is consistent with a previous study on auricular cartilage macroaggregates. 57 However, native auricular cartilage stained highly positive for VI collagen in both pericellular and territorial matrix (Supporting Information Fig. S1).
Finally, to test the feasibility of printing an auricular implant model that is easily injectable with hydrogel, we combined both PCL and alginate in one construct. There are numerous approaches to biofabrication-based auricular reconstruction, most of which use a combination of hydrogels and thermoplastic polymers. 14,15,58 Although, we also used a combination of these materials, we applied a novel approach by creating a two-part biodegradable 3D-printed mold made from PCL that is easily assembled and injectable using any type of hydrogel. Injected alginate allowed high cell survival after 21 days of culture in the auricular mold (Fig. 8). Auricular mold designs have been used previously by other groups, but were difficult to fabricate and were not integrated in the final engineered auricular tissue. 7,36,51 The novel hybrid implant model presented in this study may have several advantages over other scaffold designs: (1) PCL, which can cause inflammation when implanted in vivo, 24 will not be present in the maturing tissue, which in this design is found inside the mold, (2) the mold design can have very high mechanical properties without affecting the mechanical properties of the maturing tissue, (3) the mold design can be easily retrieved and disassembled after in vivo maturation 59 as the mold offers a partial barrier between the maturing tissue and native tissue, and (4) if necessary, the properties of the PCL mold could be adjusted to decrease the degradation rate, which can normally take up to 24 months. 24 Making use of these advantages, the mold design could then be used as follows: (1) mold assembly, (2) in vivo implantation, (3) tissue maturation, (4) retrieval of the implant model, (5) disassembly of the outer mold, (6) reconstruction using the matured cartilage tissue. Although, the auricular implant model in this study was injected with chondrocyte-laden alginate, the mold construct could be injected with any other type of hydrogel and cell source, for example, with hydrogels based on decellularized cartilage tissue. [60][61][62] CONCLUSION Here, we have described the process for engineering of a hybrid auricular cartilage implant model for facial reconstruction (see Fig. 1A). The hybrid implant model consists of a porous synthetic outer mold with high mechanical properties required to overcome forces during in vivo tissue maturation, and an inner "natural" core that consists of a biomimetic environment for cartilage tissue formation. The mold can be easily printed and assembled, while the design makes it easy to inject any suitable hydrogel for tissue formation. While long-term in vivo experiments are required to test its preclinical applicability, the work presented in this study provides a possible strategy for the use of biofabricated tissue constructs in the clinic.