Although a wide variety of natural and synthetic polymers is employed in the fabrication of scaffolds for tissue engineering, most OOC porous membranes are fabricated by only a few polymers, namely: PDMS, PC, poly(ethylene terephthalate) (PET), poly(1-caprolactone) (PCL) and Polylactic acid (PLA). Moreover, the choice of the membrane material is generally driven by practical reasons rather than the properties of the polymer related to the application. The current repertoire of commercially available flat porous membranes is limited; mainly led by PC, and PET track-etched membranes produced for Transwell inserts or filter membranes. Their fabrication is based on irradiating the material to create tracks that are then etched to produce through pores with sizes ranging from 400 nm to 8 µm. The porosity of these membranes remains low to reduce the incidence of merged tracks or double pores, and the pores density is intrinsically nonuniform due to the random nature of the irradiation. Furthermore, these membranes are too thick and rigid to represent the native basal membrane. PC track-etched membranes are extensively used in OoCs since reference data from PC transwell cell culture are widely available, thus employing PC in OoCs seemed to be a natural way of transferring knowledge from the standard cell culture system. PC is a highly hydrophobic, inert, transparent, and non-biodegradable thermoplastic. It is very stiff, with Young's modulus of approximately 2-2.4 GPa which is not comparable to any body's tissues . PET is another popular polymer for commercial membranes and inserts that is used in OoCs . PET shares many properties with PC; it is also a hydrophobic, inert, transparent thermoplastic with high stiffness (Young's modulus of 2-3 GPa) . Thus, like PC, PET is unsuitable for OoCs which apply stretch stimuli to the cells by deformation of the support membrane. Among self-produced membranes, PDMS porous membranes are to date the most employed in OoCs since they join the advantages of PDMS to those of the fabrication methods. PDMS is a low-cost elastomeric polymer characterized by physiological indifference, biocompatibility, excellent resistance to biodegradation, chemical stability, excellent gas permeability, optical transparency, and a tunable Young modulus in the physiological range (360 kPa -2 MPa) . With its flexible nature and low elastic modulus, this polymer is the most suited for applying cyclic stretch to cells. Moreover, PDMS membranes are fabricated by commonly used microfabrication techniques such as spin coating and soft lithography, which allow control over pore size and pore-topore distance in the micrometer scale . Nevertheless, this polymer also features several drawbacks that are fueling the demand for new alternative materials. PDMS is highly hydrophobic and requires surface modification approaches, as oxygen plasma, to improve its cell affinity . Besides, PDMSbased membranes absorb hydrophobic molecules such as drugs and markers commonly used in cell culture in vitro assays 130 , compromising the study results. In addition, the fabrication techniques used for PDMS-based membranes and devices have limited scalability, thus hampering production at an industrially relevant scale . As an alternative to the before-mentioned biostable materials, biodegradable polymers like aliphatic polyesters such as PLA and PCL are used to fabricate porous membranes for OoCs. PLA is one of the stiffest biocompatible polymers with Young's modulus of 3-4 GPa, making it unsuitable for membranes that should be exposed to mechanical stretch. It is, however, less hydrophobic compared to the polymers being investigated for OoCs (water contact angle of 61°) . Contrarily, PCL is characterized by a lower Young modulus of 400 MPa and a glass transition temperature of -60°C 122, resulting in a flexible polymer with a rubber-like consistency at room temperature. PCL is also one of the most hydrophobic polymers treated with a water contact angle of 140° 143 . The biodegradable nature of both polymers holds the potential for functioning as a temporary membrane, to be later replaced by the ECM of the cells creating an entirely natural cell layer. Tough, it must be considered that the pores' size might change over time due to the degradation and that the acid degradation products could affect cells. To overcome the limitations of synthetic polymer membranes, natural polymers have recently been adopted to produce membranes that better mimic the composition and function of the epithelial barrier. Arik et al. described a viable new approach to construct vitrified membranes from collagen type I hydrogel, demonstrating improved biocompatibility and are subjected to biochemical degradation by collagenase II. developed two types of collagen/elastin membranes: vitrified and non-vitrified, with stiffnesses ranging from several hundred of kPa down to 1 kPa. used a free-standing extracellular matrix gel to engineer a lymphatic vessel model. Furthermore, non-autologous proteins, such as spider silk proteins, have already been applied to form thick (3-9 μm) nonporous membranes to model the retinal pigment epithelium. Recently, Tasiopoulos et al. synthesized 470 ± 110 nm thin, mechanically robust, and fibrillar nanomembranes of recombinant spider silk proteins, where they grew endothelial and smooth muscle cells in the attempt to model the vessel wall.