FIELD

The present disclosure relates to a microfluidic based method of synthesis of porous, fibrillar collagen and porous collagen-polysaccharide microparticles for versatile tissue engineering applications, and in particular porous collagen-glycosaminoglycan microparticles.

BACKGROUND

Collagen is the most abundant protein in the human body, and an extracellular matrix (ECM) component that significantly affects many human tissues in health and disease ¹ . As a result, collagen gels have been widely used in three-dimensional cell culture of millimeter-scale structures ² . However, their application to define centimeter-scale and heterotypic tissue structures, as is often required in tissue engineering and bioprinting, has faced several limitations. The minute-long timescales for pH and thermally induced gelation ³⁴ , and the limited shear thinning behavior of collagen solutions together with the weak tensile properties of the obtained gels ² are incompatible with extrusion bioprinting high aspect ratio or cm-sized tissues. To overcome these limitations, more rapidly gelling supporting biomaterials such as alginate ⁵ have been added to collagen solutions. Macromolecular crowding agents such as high molecular weight polyethylene glycol (PEG) have been introduced to accelerate collagen gelation ⁶⁷ in the form of fiber ⁶ , sheet ⁸ , and tubular ⁹ structures.

Recently, freeform reversible embedding of suspended hydrogels has been introduced where acidic collagen solution is extruded into a support bath containing jammed gelatin microparticles ¹⁰ . Although this method allows complex acellular collagen constructs to be defined, extending it to cellular structures is challenging, due in part to the acidity of the collagen solution. As a result, cells can only be added into printed void spaces after removal of the template material. While increasing the concentration of the collagen solution enhances the mechanical properties of the bioprinted structure, it also reduces the porosity of the gels, hindering molecular transport and the migration and proliferation of cells ¹¹ .

Biomaterial microparticles, so called microgels, have been extensively prepared using conventional and microfluidic formats for a variety of applications ¹² . Microparticles have been used to encapsulate ¹³ and expand adherent cells ¹⁴ ’ ¹⁵ . Recently, microparticles were introduced as promising candidates for tissue engineering and bioprinting applications ¹² . At high packing density (>65% vol), the macro-scale voids in between jammed microparticles promote macromolecular transport, and the proliferation and migration of cells ¹⁶ ’ ¹⁷ . Consequently, casted microparticle-based matrices have been shown to improve wound healing in vivo ¹⁸ . Jammed microparticles also constitute a shear-thinning bioink that improves the fidelity during extrusion bioprinting of complex, multicellular constructs ¹⁹ . Such cellular bioinks have been used for both in vitro ²⁰ and in vivo ²¹ applications. Despite such advantages, the majority of the current microparticles are prepared from synthetic polymers such as PEG, since they are straightforward to emulsify and crosslink, a capacity that yet has to be expanded to functional microparticles from ECM proteins.

SUMMARY

As noted above, fibrillar collagens are an important constituent of the extracellular matrix (ECM) in many human tissues. Collagen hydrogels have therefore been widely used to engineer tissues in vitro and in vivo. Despite the widespread use of collagen in 3D cell culture, the capacity to define spatially organized biomaterial and tissue structures faces limitations that include poor mechanical properties of collagen gels at typical concentrations (1 to 5mg/mL), and the inability to encapsulate cells and the reduced molecular transport at higher concentrations. Collagen solutions were therefore either limited to defining acellular structures or, for cellular structures, had to be combined with other supporting biomaterials.

This disclosure describes the formation of microparticles that either solely consist of collagen or of a co-precipitate with glycosaminoglycans (collagen-GAG). The flow synthesis approach is enabled by a parallelized droplet microfluidic device that allows microparticle and pore sizes to be tuned and a fibrous micro/nanostructure to be obtained. Case studies serve to illustrate different applications of collagen microparticles including cell expansion; casted modular tissues; the formation of cell spheroids; and shearthinning, cell-laden bioinks for 3D bioprinting. Densely packed microparticles form a macro-porous matrix where the structure of microparticles and the space between them forms pore sizes of several micrometers, promoting molecular transport, cell migration and proliferation. Microparticle based matrix exhibited significantly better mechanical properties compared with their thermally gelled counterpart by three-fold reduction of tissue contraction.

Finally, to demonstrate the functionality of the microparticles for tissue engineering applications, Full-thickness in vitro skin tissues were developed at the air-liquid interface. The engineered tissue expresses functional biomarkers of the human epidermis, and the collagen-based matrix mimics the morphology and composition of the ECM of the human dermis.

Thus, the present disclosure provides a parallel droplet microfluidic system which produces highly porous microparticles of fibrillar collagen or the co-precipitate of collagen and glycosaminoglycans (GAG). The droplet microfluidic system emulsifies uniformly sized droplets containing a mixture of acidic collagen solution and a buffered solution containing either a macromolecular crowding agent such as high molecular weight PEG, or a glycosaminoglycan (GAG) such as chondroitin sulphate. In the former case, the PEG crowding agent expulses water from the collagen solution, thereby inducing compaction and the self-assembly of collagen fibrils within each droplet, forming a collagen microparticle (CP) of hierarchical organization and irregular shape. Fibril formation continues while the CPs collected from the droplet microfluidic system and incubated in a phosphate buffer solution. In the latter case, mixing of the collagen and chondroitin sulphate solutions inside aqueous droplets induces coprecipitation, forming a highly porous microparticle of collagen-GAG (CG), with an irregular, non-spherical shape.

Thus, the present disclosure provides a method of synthesizing porous collagen microparticles, comprising: simultaneously flowing a liquid containing collagen molecules, and a carrier fluid each through its own dedicated microfluidic flow conduit into an entrance of a microfluidic exit flow conduit such that said liquid containing collagen molecules and said carrier fluid mix to form a self-assembling collagen phase as they flow from the entrance to the exit of the microfluidic exit flow conduit; actively or passively breaking down the self-assembling collagen phase into collagen droplets; gelling the collagen droplets into collagen microparticles; and collecting the formed collagen microparticles.

The step of simultaneously flowing may further comprise flowing a liquid mixture containing polysaccharide molecules through its own dedicated microfluidic flow conduit along while simultaneously with flowing the liquid containing collagen molecules and the carrier fluid, the polysaccharide molecules being natural polysaccharide molecules or molecules derived from natural polysaccharide, and wherein the resulting self-assembling collagen phase is a collagen-polysaccharide precipitate with bonds formed between polysaccharide molecules and collagen molecules as the phase flows from the entrance to the exit of the microfluidic exit flow conduit resulting into collagen micro particles comprising collagen and polysaccharide.

The step of simultaneously flowing may further comprise flowing a liquid mixture containing macromolecular crowding agent molecules which are inert to the collagen molecules through its own dedicated microfluidic flow conduit, flowing a liquid emulsifying agent through its own dedicated microfluidic flow conduit, flowing a liquid buffering agent through its own dedicated microfluidic flow conduit, simultaneously with flowing the liquid containing collagen molecules and the carrier fluid, and wherein a resulting self-assembling collagen phase is passively broken down into the microparticles by as the resulting self-assembling collagen phase flows from the entrance to the exit of the microfluidic exit flow conduit.

The step of simultaneously flowing may further comprises flowing a gas stream through its own dedicated microfluidic flow conduit, and wherein the collagen phase is actively broken down into the microparticles by being chopped by air bubbles as the collagen phase flows from the entrance to the exit of the microfluidic exit flow conduit. The self-assembling is occurring as the microparticles flow from the entrance to the exit of the microfluidic exit flow conduit or as they exit or once the microparticles have exited the microfluidic exit flow conduit.

The method may further comprise the step of ice templating the collagen microparticle in which the ice templating comprises the steps of washing the microparticles and cooling the microparticles from room temperature to a temperature of about -20, or about -30, or about -50°C, at a rate between about 0.1 to about 10°C/min, thawing the collagen microparticles; or directly lyophilizing the collagen microparticles. In this aspect the method may further comprise the steps of adding a cross-linker either before or after ice templating, and washing the particles to remove unreacted crosslinker. The step of adding the crosslinker may done after the thawing step or after the lyophilisation step. The method may further comprise a step of chemically and/or physically and/or enzymatically crosslinking the collagen micro particles.

The collagen microparticles may be porous fibrillar collagen microparticles.

The collected microparticles may be washed into a buffer, and optionally a phosphate buffer, and incubated in a temperature of in a range from about 30 to about 40°C and for at least 48 hours.

The collected microparticles may be washed into a buffer, and optionally a phosphate buffer, and incubated in a temperature of about 37°C

The macromolecular crowding agent molecules may be high molecular weight amphiphilic or hydrophilic polymer molecules such as polyethylene glycol, ficoll, dextran, and polyvinyl alcohol.

The macromolecular crowding agent molecules may be selected from the group consisting of polyethylene glycol, ficoll, dextran, polyvinyl alcohol and any combination thereof.

The high molecular weight amphiphilic or hydrophilic polymer molecules may have a structure which is any one of linear, star and branched.

The macromolecular crowding agent molecules may be polyethylene glycol molecules with molecular weight ranging from about 300g/mol to about 10,000,000g/mol.

The macromolecular crowding agent molecules may be polyethylene glycol molecules with molecular weight of about 35,000g/mol. The method may be conducted at a temperature ranging from about 20°C to about 25°C.

The dedicated microfluidic flow conduits and the microfluidic exit flow conduit may have diameters in a range from about 1 pm to about 1000pm.

The dedicated microfluidic flow conduits and the microfluidic exit flow conduit may have diameters in a range from about 10pm to about 300pm.

The liquid containing collagen molecules may be an aqueous acidic collagen solution.

The liquid containing collagen molecules may comprise an aqueous acidic solution of fibril forming collagens of collagen type I, II, III, IV, V, or XI or any combination thereof.

The liquid containing collagen molecules may be an aqueous acidic solution of collagen type I.

The collected collagen microparticles may have diameters in a range from about 1 pm to about 500pm and, optionally, wherein said collected collagen microparticles have diameters in a range from about 30pm to 200pm.

The collected collagen microparticles may have a polydispersity of less than 15%.

The collagen microparticles may be formed of self-assembled collagen fibers with fiber diameters ranging between about 30nm to about 500nm, wherein the fibers show a periodic D-banding structure with about 67nm wavelength.

The collagen microparticles may be formed of self-assembled collagen fibers with a porous structure, wherein the structure has pore sizes ranging between about 1 nm to about 20pm.

The method may further comprise the step of chemically crosslinking the collagen molecules.

The present disclosure provides a collagen-based cell culture substrate comprising the porous collagen microparticles produced as disclosed above, wherein the collagen microparticles are seeded with biological cells and cultured under conditions to induce replication of the cells. The biological cells comprise any one of stem cells, primary human or animal derived cells such as epithelial and fibroblast cells, and immortalized human or animal derived cell lines. The present disclosure provides a collagen-based porous matrix comprised of a collection of collagen-based microparticles produced by the method disclosed above, wherein the collection of microparticles form a large size porous and fibrillar construct for tissue engineering, wherein the pore size may range between about 1 nm to about 200pm.

The present disclosure provides a cell-laden collagen-based porous matrix comprised of a collection of collagen-based microparticles, produced as disclosed above, and biological cells, wherein the collection of microparticles and biological cells form a large size cell-laden porous construct for tissue engineering.

A spheroid or organoid may be made by co-seeding the biological cells and the collagen microparticles.

The present disclosure provides an in-vitro functional tissue, wherein the tissue is formed of an extracellular matrix produced from the collagen-based porous collagen microparticles produced as disclosed above, and wherein the tissue-specific cells are cultured and differentiated in an interior of said extracellular matrix or on a surface of the extracellular matrix.

The present disclosure provides an in-vitro functional skin tissue comprising the collagen microparticles produced as disclosed above with the collagen microparticles acting as an extracellular matrix, and including dermal fibroblast cells and keratinocytes cells. The dermal fibroblast cells are cultured in between the collagen microparticles and the keratinocytes are cultured and differentiated over the collagen microparticles. This in-vitro functional skin tissue may further comprise fibrin formed from fibrinogen and thrombin interaction, the fibrin acting as an adhesive holding the microparticles together.

The present disclosure provides a granular cell-laden or acellular collagen-based bioink comprised of a collection of collagen-based microparticles produced by the method disclosed above, wherein the collection of microparticles with or without biological cells form a printable or shear-thinning bioink for extrusion based bioprinting.

The bioink can be extruded as cell-laden fibers with diameters ranging between about 50pm to about 1000pm. Alternatively, the bioink can be extruded as cell-laden sheet with thicknesses ranging between about 200pm to about 2000pm and with a width ranging between about 0.2mm to about 25mm. The present disclosure also provides a method of synthesizing composite porous, collagen/polysaccharides microparticles, comprising: simultaneously flowing a liquid containing collagen molecules, a liquid containing polysaccharide molecules, the polysaccharide molecules being a natural polysaccharide molecules or molecules derived from natural polysaccharide, a liquid emulsifying agent, and a liquid buffering agent, each through its own dedicated microfluidic flow conduit into an entrance of a microfluidic exit flow conduit such that said liquid emulsifying agent, collagen, polysaccharide molecules and buffering agent mix to form self-assembling composite collagen- polysaccharide microparticles with bonds formed between polysaccharide molecules and collagen molecules as they flow from the entrance to the exit of the microfluidic exit flow conduit; and collecting the composite porous collagen/polysaccharide microparticles.

The collagen/polysaccharide microparticles may be porous collagen/polysaccharide microparticles.

The natural polysaccharide may be those occurring in mammalian extracellular matrix and is known to interact with collagen.

The method may be conducted at about room temperature.

The dedicated microfluidic flow conduits and said microfluidic exit flow conduit may have diameters in a range from about 1 pm to about 1000pm.

The method collected collagen microparticles may have a polydispersity of less than 15%.

The method may further comprise the step of chemically crosslinking the collagen molecules.

The dedicated microfluidic flow conduits and the microfluidic exit flow conduit may have diameters in a range from about 10pm to about 300pm.

The liquid containing collagen molecules may be an aqueous acidic collagen solution.

The liquid containing collagen molecules may comprise an aqueous acidic solution of fibril forming collagens of collagen type I, II, III, IV, V, or XI or any combination thereof.

The liquid containing collagen molecules may be an aqueous acidic solution of collagen type I. The polysaccharide molecules may be glycosaminoglycan molecules to give collagen/glycosaminoglycan microparticles, and the glycosaminoglycan molecules may be chondroitin sulphate molecules.

The collected collagen/glycosaminoglycan microparticles may have diameters in a range from about 1 pm to about 500pm.

The collagen/glycosaminoglycan microparticles may be formed of coprecipitated collagen and glycosaminoglycan fibers with diameters ranging between about 1 nm to about 500nm.

The collagen/glycosaminoglycan microparticles may be formed of coprecipitated collagen and glycosaminoglycan fibers with a porous structure, wherein the structure has pore sizes ranging between about 1nm to about 20pm.

The present disclosure provides a collagen/glycosaminoglycan-based cell culture substrate comprising the porous collagen/glycosaminoglycan microparticles disclosed above, wherein the collagen/glycosaminoglycan microparticles may be seeded with biological cells and cultured under conditions to induce replication of the cells. The biological cells may comprise any one of stem cells, primary human or animal derived cells such as epithelial and fibroblast cells, and immortalized human or animal derived cell lines.

The present disclosure provides a collagen/glycosaminoglycan-based porous matrix comprised of a collection of collagen/glycosaminoglycan microparticles produced as disclosed above, wherein the collection of microparticles form a large size porous construct for tissue engineering, wherein the pore size ranges between about 1 nm to about 200pm.

The present disclosure provides a cell-laden collagen/glycosaminoglycan-based porous matrix comprised of a collection of collagen/glycosaminoglycan microparticles produced as disclosed above and biological cells, wherein the collection of microparticles and biological cells form a large size cell-laden porous construct for tissue engineering.

The present disclosure provides a in-vitro functional tissue, wherein the tissue is formed of an extracellular matrix produced from the collagen/glycosaminoglycan microparticles as disclosed above, and wherein the tissue-specific cells are cultured and differentiated in an interior of said extracellular matrix or on a surface of the extracellular matrix. The present disclosure also provides an in-vitro functional skin tissue, comprising the collagen/glycosaminoglycan microparticles produced as disclosed above, with the microparticles acting as an extracellular matrix, and including rmal fibroblast cells and keratinocytes cells, wherein, the dermal fibroblast cells are cultured in between the microparticles and the keratinocytes are cultured and differentiated over the microparticles. This in-vitro functional skin tissue may further comprise fibrin formed from fibrinogen and thrombin interaction, the fibrin acting as an adhesive holding the microparticles together.

The present disclosure provides a granular cell-laden or acellular collagen/glycosaminoglycan-based bioink comprised of a collection of the collagen/glycosaminoglycan microparticles produced by the method disclosed above, wherein the collection of collagen/glycosaminoglycan microparticles, with or without biological cells, form a printable or shear-thinning bioink for extrusion based bioprinting. This bioink may be extruded as cell-laden fibers with diameters ranging between about 50pm to about 1000pm.

Alternatively, the bioink can be extruded as cell-laden sheet with thicknesses ranging between about 200pm to about 2000pm and with a width ranging between about 0.2mm to about 25mm.

The present disclosure provides a collagen-based substrate comprising porous fibrillar collagen microparticles having collagen molecules selfassembled into collagen fibers with a 67nm D-periodicity, the microparticles having a polydispersity of less than 15% and fiber diameters ranging between about 1 nm to about 500 nm.

The collagen-based substrate may further include polysaccharide molecules and wherein the collagen microparticles are made of co-precipitated collagen and polysaccharide fibers, the polysaccharide being a natural polysaccharide or derived from a natural polysaccharide. The natural polysaccharide may occur in mammalian extra-cellular matrix and is known to interact with collagen. In an embodiment the natural polysaccharide is glycosaminoglycan molecules, and these glycosaminoglycan molecules may be chondroitin sulphate molecules.

The collagen or collagen/glycosaminoglycan molecules may be chemically cross-linked. io The collagen microparticles may have diameters ranging between about 1 pm to about 500pm.

The microparticles may have pores having diameters ranging between about 1 nm to about 20pm.

The microparticles may be seeded with biological cells, which may comprise any one of stem cells, primary human or animal derived cells such as epithelial and fibroblast cells, and immortalized human or animal derived cell lines.

The present disclosure also provides a method of synthesizing porous collagen microparticles, comprising: simultaneously flowing a liquid containing collagen molecules, a liquid containing macromolecular crowding agent molecules which are inert to the collagen molecules, a liquid emulsifying agent, and a liquid buffering agent each through its own dedicated microfluidic flow conduit into an entrance of a microfluidic exit flow conduit such that said liquid emulsifying agent, collagen, macromolecular crowding agent molecules and buffering agent mix to form selfassembling collagen microparticles as they flow from the entrance to the exit of the microfluidic exit flow conduit; and collecting the collagen microparticles.

A further understanding of the functional and advantageous aspects of the disclosure can be realized by reference to the following detailed description and drawings.

BRIEF DESCRIPTION OF DRAWINGS

Embodiments will now be described, by way of example only, with reference to the drawings, in which:

FIGS. 1A to 1 D show the flow synthesis of collagen-based microparticles in which:

FIG. 1A shows the architecture of microfluidic device used for flow synthesis comprised of 15 flow focusing units in which a single flow focusing unit is shown at the top, with the enlarged detail showing where the various constituents mix as they enter the exit flow conduit, a perspective view of the assembled device is shown in the middle and a top view of the assembled device is shown at the bottom;

FIG. 1B, shows a magnified image of one flow focusing unit, downstream channel, and collection channel, in which the scale bars are 200pm;

FIG. 1C shows bright-field images of the CP and CG microparticles under hydrated and critical-point dried conditions; and

FIG. 1D shows the size of collagen-based microparticles and their polydispersity versus the combined flow rate of collagen and PEG solutions in which the scale bars are 100pm.

FIGS. 2A to 2D show micro- and nano-scale characterization of the collagen-based microparticles, in which:

FIG. 2A shows SEM images of CP and CG microparticles in which the scale bars are 50pm; and

FIG. 2B shows TEM images of CP and CG microparticles in which the scale bars are 1 pm; and

FIG. 2C shows SEM images of C4P10 microparticles (collagen 4mg/mL and PEG 10%), C2P4 microparticles (collagen 2mg/mL and PEG 4%), and thermally gelled collagen (4 mg/mL), and their respective quantified graph of fiber size range in which the scale bars are 1 pm.

FIG. 2D shows TEM images and guantified fibril density and collagen area coverage of C4P10 (4mg/mL collagen and 10% PEG), C2P4 (2mg/mL collagen and 4% PEG), and thermally gelled 4mg/mL collagen in which the scale bars are 1 pm.

FIGS. 3A to 3F show studies for the application of collagen-based microparticles for cell culture and tissue engineering applications, in which:

FIG. 3A shows L/D image and viability for CP and CG microparticles and complete bio-ink, in which the scale bars are 50pm; FIG. 3B shows contraction analysis on 10mm square constructs of CP and CG microparticles and 3.5mg/mL thermally gelled collagen solution containing 1 x10 ⁶ cells/mL encapsulated ucMSCs, in which the scale bars are 5000pm;

FIG. 3C shows morphology of cells within CP and CG microparticle constructs, as well as thermally gelled collagen layer at day 5, in which the scale bars are 200pm;

FIG. 3D shows overlayed brightfield/fluorescent image of single microparticle with adherent hDFB cells at day 2 and day 10. Scale bar, 20pm;

FIG. 3E shows a fluorescent image of a spheroid formed from CP microparticles and hDFB cells, in which the scale bars are 200pm; and

FIG. 3F shows rheological analysis and brightfield/fluorescence images of extrusion bio-printed filament from CG microparticle bioink, in which the scale bars are 500pm.

FIGS. 4A to 4C show the formation of functional skin tissue using CG microparticles and hDFB and KC cells, in which

FIG. 4A shows hematoxylin and eosin (H&E) image of the in vitro tissue, in which the scale bars are 25pm;

FIG. 4B shows immunofluorescent images of the in vitro and native human tissues stained with Collagen I and KRT-10, in which the scale bars are 25pm; and

FIG. 4C shows immunofluorescent images of the in vitro tissue stained with KRT-16, Filaggring and Involucrin, in which the scale bars are 25pm.

FIGS. 5A to 5D show the flow synthesis of collagen-based microparticles in which:

FIG. 5A shows the architecture of a microfluidic device for flow synthesis comprised of 10 flow focusing units in which a single flow focusing unit is shown at the top, with the enlarged detail showing where the active breakup of collagen solution happens with the three different layers (middle left), cross- section view (Section A-A) shows the change in thickness of the microchannels for the three different layer (middle right) and a top view of the assembled device is shown at the bottom;

FIG. 5B shows the assembled microfluidic device under operation (kept on a microscope stage) on top left, a magnified image of three (out of 10) flow focusing units working in parallel on the top right, and the time-lapse shows the different stages that lead to collagen (Col) droplet formation at the bottom; and

FIG. 5C shows that collagen droplet size can be controlled by changing the air phase pressure/flow rate alone;

FIG. 5D shows the air-bubble induced breakup of collagen solution into aqueous droplets at the exit the collection microchannel (left), using a T- connector the collagen droplets are mixed with oil that caries an oil dissolvable base to neutralize collagen and initiate physical crosslinking at 37°C (right); and

FIG. 5E shows the collected and washed collagen microgels at the top and their respective size distribution and coefficient of variation at the bottom.

FIGS. 6A to 6C show the electron micrographs of non templated and ice templated microparticles along with mechanical characterization

FIG. 6A shows the particle library obtained by varying the initial collagen concentration and the freezing rates used for ice templating, top and middle row. Different pore sizes can result within a single particle; and

FIG. 6B shows that physical crosslinking of the particles enables D- periodic banding in the collagen fibers (non templated particles, top left), wall thickness varies inversely with the freezing rate of the ice templated particles (top right and middle) and the average wall thickness values are reported at the bottom;

FIG. 6C shows particles being measured using an atomic force microscope cantilever (left) and obtained Young’s modulus values for non templated, ice templated particles and ice templated particles that have undergone dehydrothermal treatment (DHT) and glutaraldehyde (GTA) based chemical crosslinking. FIGS. 7A to 7C show studies for the application of collagen-based microparticles for cell culture and tissue engineering applications, in which:

FIG. 7A shows contraction analysis on 10mm square constructs of non templated (top row), non templated crosslinked with GTA (middle row) and 1 °C/min (bottom row) microparticles supplemented with fibrinogen and 1 xio ⁶ cells/mL encapsulated ucMSCs;

FIG. 7B shows L/D image and viability for non templated particles (left), non templated particles crosslinked with GTA (middle) and 1 °C/min ice templated particles (right); and

FIG. 7C shows morphology of cells within non templated and 1 °C/min ice templated microparticle based constructs and long-term structural preservation of particles at week 3 of culture.

DETAILED DESCRIPTION

Various embodiments and aspects of the disclosure will be described with reference to details discussed below. The following description and drawings are illustrative of the disclosure and are not to be construed as limiting the disclosure. The Figures are not to scale. Numerous specific details are described to provide a thorough understanding of various embodiments of the present disclosure. However, in certain instances, well-known or conventional details are not described in order to provide a concise discussion of embodiments of the present disclosure.

As used herein, the terms, “comprises” and “comprising” are to be construed as being inclusive and open ended, and not exclusive. Specifically, when used in the specification and claims, the terms, “comprises” and “comprising” and variations thereof mean the specified features, steps or components are included. These terms are not to be interpreted to exclude the presence of other features, steps or components. As used herein, the term exemplary means serving as an example, instance, or illustration,” and should not be construed as preferred or advantageous over other configurations disclosed herein.

As used herein, the terms “about” and “approximately” are meant to cover variations that may exist in the upper and lower limits of the ranges of values, such as variations in properties, parameters, and dimensions. In one non-limiting example, the terms “about” and “approximately” mean plus or minus 10 percent or less.

The liquid containing collagen molecules may comprise an aqueous acidic solution of fibril forming collagens of, but not limited to, collagen type I, II, III, IV, V, or XI or any combination thereof. The skilled person will be aware of the potential use of one or many of the numerous types of collagen ^{31 32} . The liquid containing collagen molecules may be an aqueous acidic solution of collagen derived from vertebrates such as piscine, poultry, mammal, e.g., not limited to bovine, human, porcine, ovine, or recombinant sources. According to an embodiment, the liquid containing collagen molecules is an aqueous acidic solution of collagen type I.

Unless defined otherwise, all technical and scientific terms used herein are intended to have the same meaning as commonly understood to one of ordinary skill in the art.

The parallel microfluidic device shown generally at 10 in FIG. 1A is designed for the flow synthesis of the collagen-based microparticles and includes a plurality of flow focussing generators 11 (see top of FIG. 1A). Each of the fifteen flow focusing generators (FFGs) 11 shown in the bottom panel of FIG. 1A is supplied with 5 microchannels upstream of an 80pm-wide square orifice. The two outermost channels 14 and 16 carry equal flow rates of the continuous phase, heavy mineral oil from the oil feed microchannel 32 through a thru-hole 12. The two immediately adjacent channels 22 and 24 carry the collagen solution from the collagen feed microchannel 34 through a thru-hole 18, and the PEG or GAG solutions from the PEG or GAG feed microchannel 36 through a thru-hole 20, respectively. The middle channel 30 carries the separating buffer solution from buffer feed microchannel 38 through a thru-hole 28 that prevents unwanted mixing of the two solutions until after droplet breakup occurs once the liquids mix passes the orifice 29 at the entrance to the exit flow conduit 40.

To facilitate droplet breakup of the viscoelastic solutions at the orifice, the outlet channel 40 expands 3 times laterally and 2 times vertically compared with the orifice size 29. The fifteen (15) FFGs 11 shown in FIG. 1A ensure a sufficient throughput of the device (about 4mL/hour precursor solutions). Individual FFGs are connected to respective feed microchannels using 1 mm thru-holes. The depth of the feed microchannels (500pm) significantly exceeds the one of the orifice (80pm) to provide a uniform flow distribution between different FFGs. All individual outlet microchannels are connected via thru holes 31 to a common collection channel 41. While FIG. 1A shows fifteen (15) FFGs 11 , it will be appreciated that the present disclosure is not restricted to fifteen, and there may be more or less depending on the desired yield of microparticles.

According to an embodiment, the oil acts as a carrier fluid responsible for transporting the collagen phase, collagen droplets and/or collagen microparticles.

While the present process of synthesis is being illustrated using PEG or GAG, it will be appreciated that this disclosure is not restricted to PEG or GAG. PEG alternatives may be any high molecular weight (linear, star or branched), amphiphilic or hydrophilic polymers which are also inert (no hydrogen bonding or ionic interactions between the collagen and the PEG alternatives). Similarly, chondroitin sulphate alternatives may be any GAG or natural polysaccharides that are found in the extra-cellular matrix and known to interact with collagen, alternatively, they may be derived from such natural polysaccharides.

Collagen-based microparticles form as the emulsified droplets progress from the entrance to the exit flow conduit 40 down to the exit of the exit flow conduit 40 where the microparticles are collected. The pure collagen microparticles are formed from the molecular crowding and self assembly of collagen molecules in the emulsified droplets of collagen and PEG mix in the oil stream. Initially there is chaotic advection distributing collagen and PEG within the aqueous droplet. Then a folded/corrugated, lamella-like interconnected collagen structure is formed within the droplet. The collagen-GAG microparticles are formed from the co-precipitation of collagen and GAG molecules in the emulsified droplets of collagen and GAG mix in the oil stream.

While heavy mineral oil was used in the present studies, it will be appreciated that any hydrophobic liquid immiscible with aqueous solutions which is also inert (not reacting or forming any bond) to collagen, GAGs and other proteins and polysaccharides, such as mineral oils (heavy or light), paraffin oil, silicone oil, fluorinated oil etc.

In addition, the buffer may be any aqueous solution which is inert (not reacting or forming any bond) to collagen, GAGs and other proteins and polysaccharides, such as water phosphate buffer, Tris buffer etc.

FIG. 1B shows a bright field image of an FFG during the emulsification of collagen and PEG solutions for the formation of pure collagen microparticles (CP). The turbidity of the solution inside the emulsified droplet indicates macromolecular crowding and fibril logenesis initiate immediately downstream of the breakup location and continue while passing through the outlet channel as well as the collection channel, which takes 1 to 5s downstream of the breakup location. The experimental approach of forming collagen-GAG (CG) microparticles is similar to the described one, where the GAG solution substitutes the PEG solution. The collagen-based microparticles prepared according to one of the two approaches are collected in a 50mL centrifuge flow conduit which is pre-filled with 10mL of the fiber-formation buffer. The geometry of the FFGs is optimized for passive droplet breakup to consistently occur without the need for a surfactant to be added to the continuous phase. The droplets collected in the centrifuge flow conduit coalesce and the microparticles sediment in the buffer without the need for any washing steps, while the stratified oil phase floats atop the buffer due to its lower density.

According to an embodiment, the buffer may be a phosphate buffer ³³ ’ ³⁴ . Such phosphate buffer may act as a fiber formation buffer.

FIG. 1C shows bright field images of the CP and CG microparticles in the hydrated and critical-point dried forms. In both cases, the microparticles have a highly irregular morphology which is attributed to chaotic advection in combination with the crowding or co-precipitation processes.

Despite the irregular shape of the microparticles, their size can be controlled by the parameters of the flow synthesis process, including microchannel dimensions and flow rates. To demonstrate control over microparticle size, the flow rates of the solutions constituting the disperse phase was varied while maintaining the oil flow rate constant at 120pL/min. FIG. 1D shows representative micrographs and the extracted microparticle size at different flow rates. The indicated flow rate is the combination of the collagen and PEG solution flow rates while the buffer flow rate was kept at either 10 or 20pL/min, depending on the breakup regime. At low combined flow rates, 20 - 40pL/min, the breakup occurs within the orifice resulting in the dripping regime, forming CP microparticles in the range of about 25 to about 70pm. Increasing the flow rates further postpones the breakup into the outlet channel, corresponding to the jetting regime and larger microparticles of about 100 to about 200pm. The generated microparticles at each flow rate have a polydispersity of about 15% or less, which provides control over the size of the microparticles. Such polydispersity allows for the selection of the proper microparticle size depending on the desired application.

To further analyze the morphology and nanostructure of the samples, scanning and transmission electron microscopy (SEM and TEM) were performed on the CP and CG microparticles, as shown in FIG. 2A and 2B. The SEM image of the CP sample exhibits a fibrous morphology, whereas the CG microgels have a folded-sheet like morphology (FIG. 2A). TEM images taken from 70nm thick slices serve to characterize the nanostructure of the particles (FIG. 2B). The 11 ,5kx TEM images show both microparticles exhibit a highly porous structure with pore sizes extending from the nanoscale up to several micrometers. The folded-sheet morphology of the CG microparticles is also seen in the TEM image. The 50kx TEM image of the CP sample shows the D- banding pattern with 67nm wavelength within collagen fibers/filaments, confirming the physiologically relevant self-assembly of collagen fibrils. The CG microparticles have a different morphology at 50kx magnification, where the folded sheets consist of a dense network of 30 - 80nm long fibers with diameters of 5 - 10 nm. Based on the SEM and TEM images, both microparticle types include a highly porous structure, which is crucial for molecular transport and cell proliferation and migration.

The described flow synthesis route allows tuning the collagen density and porosity, the mechanical properties, and the nanostructure via controlling the concentration of the precursor solutions. To demonstrate this capability and the superiority of the collagen based microparticles over thermally gelled collagen constructs, CP microparticles were synthesized with collagen and PEG concentrations of 2mg/mL and 4%, respectively (C2P4), as well as 4mg/mL and 10% (C4P10), along with thermally gelled collagen construct at 4mg/mL concentration as the control group. As shown in FIG. 2C, the thermally gelled collagen construct consists of collagen fibers in the range of 30nm - 90nm, while the fiber size in CP microparticles is significantly larger (up to about 250nm for C2P4 and 360nm for C4P10) due to the molecular compaction by PEG. The D-banding structure is also clearly visible in the collagen fibers in the C4P10 and C2P4 samples. The TEM images of the samples shown in FIG. 2D also depict that in the thermally gelled collagen construct, only less than 3% of the area is covered by collagen, while for the C2P4 and C4P10 samples the covered areas are about 14% and 27%, respectively. Such control over the fiber size and fibril density enables modulation of the mechanical properties of the collagen based microparticles.

Case studies serve to illustrate the utility of collagen-based microparticles for versatile applications in cell culture and tissue engineering. To establish cytocompatibility, primary human dermal fibroblasts (hDFB) and umbilical cord derived mesenchymal stromal/stem cells (ucMSC) were cultured with collagen-based microparticles. Cell viability was assessed with a fluorescent Live/Dead kit. hDFBs were directly seeded on the CP and CG microparticles, while the ucMSCs were used along with CP microparticles in a 10mg/mL fibrinogen solution to form a bioink, crosslinked with a 50U/mL thrombin solution. The viability analysis and a representative micrograph of hDFB on collagen-based microparticles are shown in FIG. 3A. In all cases the cells exhibited >90% viability after 5 days of culture. An advantage associated with preparing cell-laden 3D matrices from the collagen-based microparticles presented in this study over the casting layers of similar thickness from thermally gelled collagen solution is the improvement in mechanical properties. For CP microparticles, the improved mechanical properties are due to the local compaction associated with macromolecular crowding ^{6 8} , while for CG microparticles, it is due to coprecipitation with chondroitin sulphate ²² . To compare the improved mechanical properties of the matrix made of collagen-based microparticles with their counterparts made from a casted thermally gelled collagen layer, contraction of their cell-laden square construct was monitored over 5 days of culture. A solution of 3.5mg/mL collagen was used for the formation of both CP and CG microparticles, as well as the thermally gelled collagen constructs. Using custom-made poly(dimethylsiloxane) (PDMS) molds, square constructs of 10mm x 10mm with 1 mm thickness were made using the call-laden CP, CG, and the collagen solution, where ucMSCs were included in the three cases at a final density of 1 xio ⁶ cells/mL. The three samples were cultured for 5 days, and images were recorded every 24 hours to compare the contraction of the square constructs over the culture period.

The sample images of the construct at days 0, 2 and 5, along with their graph are shown in FIG. 3B. The samples of the thermally gelled collagen exhibited the highest level of contraction, shrinking to a spherical shape at day 5, while the two microparticle based constructs maintained their square shape with a size that reduced during culture. Comparing the two microparticle-based samples also indicated the CG sample to contract less, where its size at day 5 is 20% larger than its CP counterpart. The difference with the improved mechanical properties was associated to co-precipitation with chondroitin sulphate.

The cell morphology for the three cases were also studied at day 5, where the cells were fixed and the F-actin filaments of their cytoskeleton were stained, as illustrated in FIG. 3C. It can be observed that in the thermally gelled construct, a large number of cells is locally elongated along the same direction due to shrinkage, whereas in both CP and CG microparticle constructs, the cells have proliferated and formed a cellular network in the space between the microparticles. This allows for the longer-term cell culture for tissue engineering applications, for example in vitro tissue engineering application.

Since the collagen-based microparticles presented in this study are cyto- compatible with the available cell binding sites on their collagen component, they may be used as a cell culture substrate whether for suspension-based cell culture, or as a micro-carrier for in vivo cell delivery. In a case study that serves to illustrate the capacity of using collagen-based microparticles as cell culture substrates, hDFBs were seeded with a low density of the CP microparticles in a non-adherent culture plate and cultured for 10 days. Sample images of the microparticles with the cells attached on their surface are shown in FIG. 3D. It can be observed that while at day 2 a small surface of the collagen-based microparticles is covered with the cells, after 10 days the cells have proliferated and covered the surface of the microparticle.

Cell spheroids and stem cell derived organoids are multicellular structures that provide physiologically relevant structural and functional properties ²³ . The presented collagen-based microparticles presented here have the potential to support the formation of such cell spheroids and organoids. The suspension of cells and microparticles may be used with the commercial platforms ²⁴ , or simply pipetted in a conical bottom microtube, cultured for 12 to 24 hr until the cells and microparticles aggregate, and then transferred to a nonadherent culture plate. As a case study that serves to demonstrate this capability, hDFB spheroids were formed using CP microparticles and cultured for 10 days. FIG. 3E shows the immunofluorescence image of a representative spheroid, where the F-actin filaments of the cells are stained.

The microparticles developed in this study have a great potential for bioprinting applications. Suspending the collagen-based microparticles at high packing density (>65%vol) along with cells constitutes a shear-thinning bioink candidate for extrusion based bioprinting. To demonstrate such capability, 1 M cells/mL of ucMSCs was suspended with a high concentration of collagen- based microparticles, centrifuged at 500G for 2min, followed by removing the phase-separated culture media from the compacted microparticles. The bioink was then gently mixed manually using a syringe to homogenize the cell distribution. Rheological analysis was performed on the bioink, and the results are shown in FIG. 3F. At low strain rates the elastic properties of the bioink dominate. With increasing the strain rate, the storage modulus drops while the loss modulus increases, increasing the ‘fluidity’ of the bioink. Such behavior is desirable for extruding the bioink through a printhead or a needle. The recovery analysis reveals that by alternating the shear rate between 0.5% and 500% at 100s time intervals, the dominance of the storage and loss moduli alternate. At a low shear rate of 0.5%, the storage modulus is highly dominant, exhibiting solid-like properties and maintaining the structure of the bioink. Switching to a high shear rate of 500%, the storage modulus decreases significantly, and the loss modulus becomes dominant, providing the required fluidity to the modular bioink and allowing for extrusion.

Once the shear rate is switched back to 0.5%, the bioink recovers its solid-like properties as the storage modulus retains its original value. The presented rheological behavior is desirable for favorable ‘printability’ of a bioink, where the bioink can be extrusion bioprinted and rapidly retains it structural integrity thereafter. Filaments were extruded using a 16G needle and cultured for 5 days in a non-adherent culture plate. FIG. 3F shows bright field and immunofluorescence (F-actin) images of the filament after 5 days of culture.

The capacity of the collagen-based micreoparticles to promote the formation of functional human tissues was demonstrated for the case study of a full thickness skin tissue, cultured at an air-liquid interface (ALI). The CG microparticles were selected as coprecipitates of collagen and GAGs as they are major components of the dermal ECM and have shown great promise in regenerating human skin tissues ²² . Therefore, CG microparticles were selected to recapitulate the ECM of the human dermis. To develop in vitro human skin tissues, initially CG microparticles and 5*10 ⁵ cells/mL hDFBs were cultured at 0.5mm thickness for 2-3 days to form a planar construct. Then human primary keratinocytes (KCs) were seeded on their surface at 5X10 ⁵ cells/cm ² and cultured for 3-4 day, until the KC layer was fully confluent. Circular constructs were then gently placed onto track-etched polycarbonate membranes (pore size 3pm) of 12-well inserts, cultured at ALI for 4 to 5 weeks, fixed, paraffin embedded, and sectioned for immunofluohistochemistry. The hematoxylin and eosin (H&E) and immunofluorescence (IF) images obtained from the in vitro skin tissue are shown in FIGS. 4A to 4C. The H&E image (FIG. 4A) reveals the bi-layer structure of skin, with a multi-layer epidermis formed on the dermal equivalent. To evaluate the morphology of the dermal layer, immunofluorescence stains for collagen I of the in vitro tissue are compared with the ones for native human skin tissue (see FIG. 4B). It could be observed that the expression of Collagen I and the morphology of the matrix recapitulate aspects of the ECM organization in native human skin, suggesting the capacity of the CG microparticle-based structures to support the development of functional in vitro skin tissues. Differentiation of the epidermal layer was studied by analyzing the expression of the cornified envelop proteins in the stratum corneum ²⁵ . FIGS. 4B and 4C show immunofluorescence images of in vitro tissues stained for KRT-10, KRT-16, Filaggrin, and Involucrin, and display positive marker expression in the epidermal layer, as well as the formation of a differentiated epidermis.

Conclusion

Collagen is the most abundant protein in the ECM of several human tissues, particularly skin, and plays a crucial role in tissue engineering strategies. However, its application for the formation of casted and bioprinted tissues faces multiple challenges. Thermal gelation of neutral pH collagen solution requires several minutes which either significantly limits the structural complexity or the rate of formation of produced tissues. In addition, the typical concentration range (1 to 5mg/mL) is associated with poor mechanical properties and significant contraction of cellular constructs during remodeling. At higher concentrations, the mechanical properties improve and remodeling is reduced, however, the small pore size of the matrix is undesirable for encapsulated cells. Therefore, approaches such as freeform reversible embedding of suspended hydrogels, and the use of macromolecular crowding agents improve printability but their dependency on acidic collagen solutions prevents the incorporation of cells. Co-precipitation of collagen with glycosaminoglycans has also been extensively used in forming dermal templates for in vivo and routine clinical use. Similar to molecular crowding, this method also has been limited to casting pre-defined planar acellular constructs. To overcome these limitations and address the requirements of tissue engineering and bioprinting applications, a microfluidic flow synthesis method for collagen-based microparticles was introduced. Macro-porous fibrillar microparticles of pure collagen and collagen-GAG were formed, where the particle size and porosity were controlled by the flow synthesis parameters such as precursor flow rates. Using human dermal fibroblasts (hDFB) and umbilical cord derives mesenchymal stem cells (ucMSC), the data has shown the capacity of the presented collagen and collagen-GAG microparticles to serve as cell culture substrates, to enable casting of tissues, forming spheroids, and as shear-thinning bioinks for bioprinting. At high packing densities, jammed collagen-based microparticles along with cells constitute a shear-thinning bioink for extrusion bioprinting applications. Bioprinted constructs exhibit a high porosity matrix due to the space between the microparticles, facilitating macromolecular transport, and cell proliferation and migration. The results demonstrated for the collage-based microparticles to improve the mechanical properties of the matrix, reducing tissue contraction 3 to 4 folds in comparison with thermally gelled collagen solution.

Finally, to demonstrate the applicability of the collagen-based microparticles for forming functional human tissues, a full thickness in vitro skin tissue was cultured at the air-liquid-interface (ALI). The collagen-GAG microparticles and hDFBs were used to form planar dermal equivalents, and then KCs were cultured on their surface. The tissues were cultured at ALI for 4 - 5 weeks. The histology and immunofluorescence analysis of the in vitro skin revealed the layered skin structure. The collagen composition and morphology of the dermal matrix mimicked the one of the native human dermis. The epidermal layer expressed the functional biomarkers of the differentiated epidermis, including KRT-10, KRT-16, Filaggrin, and Involucrin.

Materials and methods

Materials

Polydimethylsiloxane (PDMS) elastomer kit was purchased from Ellsworth Adhesives Co. (Germantown, Wl, USA); SU-8 2150 and its developer were purchased from Kayaku Advanced Materials, Inc. (Westborough, MA, USA); high glucose Dulbecco’s modified eagle medium (DMEM) containing 4,500mg/L D-glucose, L-glutamine, and 110mg/L sodium pyruvate, fetal bovine serum (FBS), penicillin and streptomycin (penicillin 10,000 units/mL and streptomycin 10,000 pg/mL), live/dead staining kit, 0.25% trypsin-EDTA, Chondroitin sulfate sodium salt from shark cartilage, 35kDa PEG, fibrinogen and thrombin from bovine plasma, L-ascorbic acid, 7.5% solution of bovine serum albumin, hydrocortisone, SITE+3 liquid media supplement (100x), calcium chloride dihydrate, and Dulbecco’s phosphate buffered saline (DPBS) were purchased from Millipore Sigma (Oakville, ON, Canada). AggreWell™800 and anti-adherence rinsing solution were purchased from STEMCELL Technologies (Vancouver, BC, Canada).

Anti-Filaggrin antibody (GTX37695) was purchased from GeneTex, Inc. (Irvine, USA), Anti-Collagen type 1 (E8F4L) antibody was purchased from Cell Signaling Technology (Danvers, MA, USA), Anti-Cyto keratin 10 antibody (DE- K10) and anti-lnvolucrin antibody (SY5) were purchased from abeam (Toronto, Canada), Anti-Cytokeratin 16 Monoclonal Antibody (LL025) was purchased from Thermo Fisher Scientific (Waltham, MA, USA), the fluorescence conjugated secondary antibodies (Alexa Fluor® 488 AffiniPure Donkey AntiRabbit IgG, Alexa Fluor® 647 AffiniPure Donkey Anti-Mouse IgG), and IgG-free bovine serum albumin (BSA) were purchased from Jackson ImmunoResearch Inc. (West Grove, PA, USA). The 4’,6-Diamidino-2-Phenylindole, Dihydrochloride (DAPI), and Phalloidin-Alexa Fluor™ 488 were purchased from Thermo Fisher Scientific (Waltham, MA, USA). 10x PBS, 0.5 N sodium hydroxide, and 10 mg/mL FibriCol® bovine collagen type 1 were purchased from Advanced BioMatrix Inc. (San Diego, CA, USA). The formalin and 37% formaldehyde solutions were purchased from VWR (Mississauga, ON, Canada). The triton X-100 was purchased from Bio Basic Canada Inc. (Toronto, ON, Canada). Heavy mineral oil was purchased from Fishe Scientific (Walthman, USA). Surface hydrophobic treatment agent was purchased from Aquapel (Lachute, Canada). All other reagents were of analytical grade.

Device Fabrication The multi-layer droplet microfluidic systems shown generally at 10 in FIG. 1A were fabricated using standard photolithography and soft lithography methods. The flow distribution and flow focusing units were fabricated using two different SU-8 master molds. The mold for distribution layer was fabricated in a single step using SU-8 2150 at 500pm thickness. Briefly, SU-8 was coated on a silicon wafer using a spin coater, and then was baked at 65°C and 95°C according to the supplier’s protocols. Then the microchannels were patterned through a chromium coated glass photomask using a EVG 620 aligner, followed by post exposure baking at 65°C and 95°C.

The wafer was then developed in SU-8 developer on an orbital shaker to removed non-cured photoresist. The wafer was then rinsed thoroughly using isopropanol and dried by a nitrogen stream. To assure the durability of the wafer, it was hard baked at 180°C for 30min. The mold for flow-focusing units was fabricated at two sequential layers, each 80pm. Photomasks for patterning the two layers were aligned using the EVG 620 aligner. The first layer was coated, UV-exposed and post baked, then the second layer was coated, UV- exposed and baked, followed by developing and hard bake.

PDMS elastomer and its curing agent were mixed at 10:1 ratio and degassed. A thin layer of PDMS was poured on the flow-focusing wafer, debubbled, and cured in an oven at 80°C for 1 hour. A thin layer of PDMS was poured on the distribution wafer and de-bubbled, then the flow focusing PDMS layer which was already cured, was flipped upside, and placed over the distribution layer. The two layers were aligned and placed in the oven to cure. Next step, the two sides of the PDMS layer were connected by punching 1 mm holes using a biopsy punch. A thick PDMS block, separately prepared, was bonded on the flow focusing side of the PDMS layer after corona treatment. The 2mm inlet and outlet holes were punched using a 2mm biopsy punch, and the distribution side of the layer was bonded on a microscope slide following corona treatment. The device was placed in the oven 80°C for at least 4 hours to assure a strong bond between different layers. Afterwards, to increase the hydrophobicity of the microchannels, Aquapel was injected in the device through one of the inlets and incubated for 1 to 2min, then it was removed using a syringe and the microchannels were flushed suing mineral oil. The device was then baked in the oven at 80°C for 1 hours.

Microparticle Synthesis

Throughout this study, the flow synthesis of collagen and collagen-GAG microparticles were performed using acidic solutions made or diluted in DI water. However, it was tested and proved that the process functions at neutral pH and in presence of 1 x buffer ions (e.g., PBS or Tris buffer) as well. Acidic collagen solution was diluted in acetic acid containing water (pH 2) at the concentrations of 2, 3.5 or 4mg/mL. Chondroitin sulfate was dissolved in DI water at the concentration of 4mg/mL. PEG solution was prepared by dissolving lyophilized PEG in buffer (pH 7) at the concentration of 100mg/mL. For the flow synthesis of collagen microparticles, oil, collagen, buffer, and PEG solutions were loaded in BD syringes and connected to the microfluidic device using Tygon tubes. A Tygon tube was connected to the outlet and was directed into a centrifuge tube pre-filled with 10mL of pre-filled phosphate buffer. The syringes were connected to 4 parallel Harvard syringe pumps and ran at desired flow rates. The microfluidic system mixed the solutions and emulsified in the oil stream. The outlet tube was placed in the collection centrifuge tube such that the outlet solution containing the droplets were accumulated over the buffer solution. Since no surfactant was used in the oil, the droplets gradually merged into the initial buffer solution, separating the synthesized microparticles from oil.

The collagen microparticles were then washed into fresh phosphate buffer and incubated in a 37°C incubator for at least 48 hours to assure the selfassembly of collagen molecules. The collagen-GAG microparticles were synthesized with the same procedure, except instead of PEG solution, a solution of chondroitin sulfate was used. Collagen-GAG microparticles also did not require the incubation step in the phosphate buffer. To make phosphate buffer ⁸ , 7.89mg/mL sodium chloride, 4.26mg/mL dibasic sodium phosphate, and 10mM Tris were dissolved in DI water and the pH was adjusted to 7.4.

Electron Microscopy

For scanning electron microscopy (SEM), microparticles were crosslinked using a 2% glutaraldehyde solution for 2 hours at room temperature, then dehydrated in ethanol in sequential steps (30%, 50%, 70%, 90%, 100%x3 ethanol), 5 to 10min each. The microparticles were then dried using a BAL-TEK critical-point drier (CPD 030) and placed on the SEM pin mount, followed by coating 5 to 10nm of gold using a Leica ICE sputter coater. The samples were then imaged using ZEISS microscope.

For transmission electron microscopy (TEM), microparticles were crosslinked using a 2% glutaraldehyde solution for 2 hours at room temperature and then rinsed 3 with DI water times, each time 5 min, followed by gentle centrifugation and removal of supernatant. The microparticles were then incubated with osmium tetroxide (2% vol/vol) and potassium ferrocyanide (1.5% wt/vol) for 1 hour at room temperature, followed by 3 times washing with DI water. Next, the microparticles were incubated with 1 % (wt/vol) tannic acid for 2 hours at room temperature and washed 3 times with DI water. The microparticles were then incubated in a 1 % (wt/vol) solution of uranyl acetate at 4°C over night and then washed with DI water 3 times. In the next step, the microparticles were dehydrated in ethanol in a sequential process as explained above, and then the ethanol was replaced with acetone in a similar procedure. In the final step, the microparticles were infiltrated in Agar 100 Hard resin in a sequential diluted manner (10%, 30%, 50%, 70%, 90%, and 100%x2), each step 1 to 2 hours and in 100% Agar 100 Hard over night with gentle agitation on a orbital shaker. The sample was then placed in molds, centrifuged (100G for 30s) and cured in an oven at 60°C for 72 hours. For imaging, the samples were cut at 70nm thickness and imaged using a T20 Phillips TEM microscope.

Cell Culture

Primary human neonatal dermal fibroblasts (hDFB) isolated from human foreskin, were purchased from ATCC® (PCS-201-010™, Manassas, VA, USA). Human primary dermal epidermal Keratinocytes (H-6066) were purchased from Cell Biologies Inc. (Chicago, IL, USA). Human umbilical cord derived mesenchymal stem cells (ucMSC) were a kind gift from Dr. Jeschke lab at Sunnybrook Research Institute. hDFBs were cultured in serum free growth kit from ATCC (PCS-201-040) and used for the experiments in passage 2 or 3, as recommended by the supplier. Keratinocytes were cultured in the complete human epithelial cell medium kit from Cell Biologies (H6621) and were used for experiments between passages 5 to 15. ucMSCs were cultured in high glucose DMEM supplemented with 10% FBS and 1 % Pen/Strep and were used for experiments between passages 2 to 5. For all cell types, culture media was changed every 48 hours until they reach 80% confluence. Then they were rinsed with PBS and trypsinized for 3 to 5min to detach. Culture media was added twice the volume of the cell suspension, and then centrifuged at 300G for 5 min. The supernatant was aspirated, and the cells were resuspended in fresh culture media, followed by cell counting.

For cell culture on dispersed microparticles, a low concentration suspension of microparticles and 5*10 ⁵ cells/mL cells were seeded in a non adherent culture plate, treated with anti-adherence rinsing solution, and culture until analysis, with culture media being changed gently every 48 hours. For the formation of microparticle-based constructs, cells at 6x1 o ⁶ cells/mL were added to the compact microparticles and gently mixed with a pipette with the tip cut using a sterile blade. The high packing density suspension was them pipetted to the desired well or cast and allowed for 6 hours for the cells to attach. Culture media was then added to the wells and replaced every 24 hours until the construct was used for analysis.

For the formation of cell spheroids, diluted suspension of microparticles and 2x1 o ⁶ cells/mL were either added to AggreWell™800 or into a sterile conical centrifuge tube and allowed the spheroid to form for 24 hours. The culture in AggreWell™800 was continued until analysis, while the spheroids in centrifuge tube were transferred to a non-adherent culture plate and cultured until analysis. The bioink was formed by adding a solution of 100mg/mL fibrinogen and 6x1 o ⁶ cells/mL cell suspension to a high packing density microparticle suspension, with the volumes adjusted to reach a final concentration of 10mg/mL for fibrinogen and 1 xio ⁶ cells/mL for the cells. The suspension was then centrifuged at 300G for 1 min and a volume equivalent to the sum of fibrinogen solution and cell suspension was removed from the bioink. The bioink was then gently mixed using syringe.

Viability and Morphological Analysis For viability analysis, the constructs were incubated with a pre-made solution of Live/Dead stains for 1 to 2 hours, and washed 3 times with PBS, 5min each. The constructs were then imaged using a Nikon confocal microscope. Viability was calculated by dividing the number of live cells to the total number of live and dead cells. For morphological analysis, the cells were fixed using a buffered solution of 3.7% formaldehyde and 0.3% Triton x-100 for 30 min. They were then washed 3 times with PBS and incubated with a solution of Phal loidi n-FITC and DAPI in PBS for 2 hours at 4°C. The constructs were then washed 3 times with PBS and imaged using an Axio Observer ZEISS microscope.

Rheological Analysis

Rheology analysis was performed on the bioink using a Discovery Rheometer (DHR-3) with a 60mm cone and plate geometry. The tip gap was set to 100pm, and the temperature was set to 25°C. To measure the shearthinning properties of the bioink, shear rate was ramped from 0.1 % to 1000% at 1 kHz with 5 points every decade, and the storage and loss moduli (G' and G", respectively) were measured. For the analysis of shear recovery, 1% and 500% strains at 1 kHz were applied periodically, each strain sate for 100s and a total of 3 alternation between low and high strains.

Skin Tissue Culture

To develop in vitro skin tissues, a high packing density suspension of CG microparticles and hDFB at the density of 5*10 ⁵ cells/mL were casted in planar form and cultured for 2 to 3 days to form the dermal equivalent. Then, Keratinocytes at the density of 5x1 o ⁵ cells/cm ² were seeded over the dermal equivalents and cultured submerged using a mix of half/half dermal and epidermal culture media for 3 to 4 days, until the Keratinocytes are fully confluent. The planar constructs were then transferred into 12-well polycarbonate cell inserts with 3pmm pores and cultured at the ALI for 4 to 5 weeks. Each well was filled with 900pL of ALI culture media and replaced every 24 hours. After 5 weeks, the constructs were fixed using a buffered solution of 10% formalin overnight at 4°C and used for processing. To make 500mL of ALI culture media, 475mL of serum-free epithelial basal medium was used and the following supplements were added to it: 7.5% bovine serum albumin (6.67mL), 10 mg/mL L-ascorbic acid (2.5mL), 50 pg/mL hydrocortisone (5mL), 100x SITE+3 supplement (5mL), 1 mol/L calcium chloride (600pL), and Pen/Strep (5mL).

Tissue Processing and Immunostaining

Tissue processing and staining were performed in the STTARR Innovation Centre, University Health Network (UHN), Toronto, ON, Canada. After fixation, the tissues were placed in cassettes and placed in a Histo-Tek® VP1 ™ (SAKURA) instrument for sequential ethanol dehydration and embedding in paraffin wax. The embedded tissues were then cut in 4pm slices using a Shandon™ Finesse™ ME+ microtome (Thermo Scientific) and placed on microscope slides.

For immunofluorescence staining, the sections were deparaffinized using the protocol for each antibody, blocked using a solution of 5% goat serum and 0.3% Triton x-100 in PBS. After blocking, the sections were incubated with the antibodies diluted according to the manufacturer’s ratio in a solution of 1 % goat serum at 4°C, followed by 3 times washing and incubation with the proper conjugated secondary antibody, and then counterstaining with DAPI. The slices were then washed 3 times with PBS and mounted with mounting medium. Fluorescence microscopy was then performed using an Axio Observer ZEISS microscope.

Statistical Analysis

All data is presented as mean ± standard deviation. Statistical analysis was performed using ANOVA or t-test. All results are based on 3 or more independent experiments.

According to an embodiment, to form a bioink that may 1) better mimics the multiscale architecture of collagen in tissues, 2) allow for facile cell delivery, and 3) prevent contraction during day-long culture, a bioink consisting of macroporous collagen particles (MCP) supplemented with a temporary adhesive (fibrinogen) was systematically prepared. While the MCP internal/external porosity allows recapitulation of the microstructure of targeted tissue, the temporary adhesive holds the microgels together to allow submerged culture.

The fibrinogen may be turned over with time, and the ECM secreted by the cells may replace the fibrinogen. The cell-secreted ECM and the proliferating cells themselves may provide mechanical integrity thereafter to the construct. According to an embodiment, the fibrinogen may be turned over in about a week.

According to an embodiment, a parallelized flow focusing device was designed, fabricated, and perfused with collagen solution (tested up to 10 mg/ml), forming microgels at an about 45x higher viscosity than the highest previously reported concentration (about 2 mg/ml) ^26-30 . The collagen solution is prepared as follows; the total collagen solution (10 mg/ml) volume (VC) is determined and is mixed with 1/10 parts (VC /10) of 10X phosphate buffer solution (PBS). The solution can be diluted with acidic (pH 2) DI water while readjusting the PBS volume (keeping the same 1 :10 ratio) to attain the desired collagen concentration. To facilitate consistent breakup, an active mechanism was required which utilizes an air bubble to induce the hydrodynamically confined jet formed by the collagen solution to break up. This emulsification technique was termed “air bubble chopper” (ABC) method, a technique that can be used to emulsify high viscosity biopolymer solutions. A parallelized device with 10 flow focusing units being tested up to a throughput of up to 2.4 ml/hr. The increase in protein concentration, in this case, collagen, allows tuning of particle stiffnesses and morphology needed for optimizing the scaffold’s response to the tensile field generated by the cells.

According to an embodiment, the parallel microfluidic device shown generally at 100 in FIG. 5A is designed for the flow synthesis of the collagen- based microparticles and includes a plurality of flow focussing generators 110 (see bottom of FIG. 5A). Focussing generators 110 in FIG. 5A is the same as focussing generators 11 in FIG. 1A. Focussing generator 110 at the top of FIG. 5A is an example of one of the ten focussing generators 110 along the bottom of FIG. 5A. Each of the ten flow focusing generators (FFGs) 110 shown in the bottom panel of FIG. 5A is supplied with 4 microchannels upstream of an 80pm-wide and 160 pm-high orifice. The two outermost channels 160 and 180 carry equal flow rates of the continuous phase, heavy mineral oil with Span 80 surfactant from the oil feed microchannel 320 through a thru-hole 120. The two immediately adjacent channels 220 and 240 carry the collagen solution from the collagen feed microchannel 340 through a thru-hole 140, and the air stream feeds microchannel 380 through a thru-hole 280, respectively.

To facilitate droplet breakup of the collagen solution and air at the orifice, the outlet channel 400 expands 5 times laterally and 1 .625 times vertically compared with the orifice size 290. The ten (10) FFGs 110 shown in FIG. 5A ensures a sufficient throughput of the device (about 2.4mL/hour of collagen solution). Individual FFGs are connected to respective feed microchannels using 1 mm thru-holes. The depth of the feed microchannels (500pm) significantly exceeds the resistor section (Layer 1 ) of channels 160, 180, 220, and 240 (40pm) to provide a uniform flow distribution between different FFGs. All individual outlet microchannels are connected via through holes 310 to a common collection channel 410. While FIG. 5A shows ten (10) FFGs 110, it will be appreciated that the present disclosure is not restricted to ten, and there may be more or less depending on the desired yield of microparticles.

While the present process of synthesis is being illustrated using air, it will be appreciated that this disclosure is not restricted to air. Air alternatives may be any gas e.g., nitrogen, argon, oxygen etc.

Collagen-based microparticles form as the emulsified droplets progress from the entrance to the exit flow conduit 400 down to the exit of the exit flow conduit 400 where the microparticles are collected. The pure collagen microparticles are formed from the in-line physical crosslinking after exiting collection channel 410 outside of the device, however physical crosslinking can also be triggered within the outlet channel 400 and/or the collection channel 410. According to an embodiment, the oil acts as a carrier fluid responsible for transporting the collagen phase, collagen droplets and/or collagen microparticles.

While heavy mineral oil was used in the present studies, it will be appreciated that any hydrophobic liquid immiscible with aqueous solutions, such as mineral oils (heavy or light), paraffin oil, silicone oil, fluorinated oil etc.

According to an embodiment, the formation of the collagen microparticles may be obtained with the following steps:

1 . The collagen phase enters the orifice and begins to form a bulb.

2. This growing bulb then moves downstream of the orifice and continues to increase in size.

3. The air phase begins to protrude inside the orifice to form its own bulb.

4. The air phase’s bulb grows in size and begins to choke the orifice causing the collagen bulb to form a thin neck (inside orifice).

5. The air bubble eventually occupies the entire orifice which then perturbs/dissects the elongated collagen neck

6. A collagen droplet is formed downstream of the air bubble.

Ice templating was used to form macroporous (porosity about 95%) particles with pore sizes decreases with increasing freezing rate. The high local collagen compaction in combination with chemical crosslinking (e.g., glutaraldehyde, GTA) allowed to further tune the particle stiffnesses (3-30 kPa) and to tailor degradation profiles of the particles. Umbilical cord-derived mesenchymal stem/stromal cells (UC-MSCs, 100k cells/cm ² ) were cultured with both non-templated and ice-templated collagen microgels (10 mg/ml) supplemented with fibrinogen solution (10 mg/ml, clotted with thrombin). The construct prepared from these MCPs and cells retained their microarchitecture for > 3 weeks while the structure prepared from MCP-based bioink only compacted insignificantly (< 5%). This marks a significant improvement in comparison to the scaffolds prepared from collagen-GAG and collagen-PEG particles which reduce to 50% size in < 1 week.

According to an embodiment, to emulsify highly viscoelastic solutions such as high concentration collagen an active breakup technique is deployed to collapse the thread at the orifice section of the flow focusing device (at the beginning of Layer 3, FIG. 5A). At the same time to ensure high emulsification throughput and a wide droplet size range, 10 flow focusing units are run in parallel with each unit having three separate layers (FIG. 5A). Layer 1 is a resistance adding layer which ensures uniform flow distribution to all 10 units from the respective distribution microchannels. While Layer 2 uses thickness values 1 x target diameter of collagen droplets, Layer 3 provides more space to produce larger droplets and further improves droplet breakup performance. The distribution layer and the collection microchannels are connected to each other via through-holes.

The assembled device (FIG. 5B) can be operated at room temperature without any premature gelation or device failure. As seen, even the high expansion ratio from Layer 2 to Layer 3 does not passively break up due to a hydrodynamic instability at high collagen concentrations (e.g., the about 10 mg/ml selected here). An additional air stream was supplied to the flow focusing unit to induce passive breakup of an air bubble to trigger the breakup of the hydrodynamically focused collagen solution downstream of the orifice region of the flow focusing unit. In the absence of air supply, the collagen solution forms an elongated filament/jet downstream of the orifice and fails to passively breakup into droplets. However, when air is introduced, the collagen solution protrudes to form a rounded projection shaped as a bulb downstream of the orifice and behind the recently produced air bubble. The breakup of collagen solution happens as follows; the collagen phase enters the orifice and begins to form a bulb. This growing bulb then moves downstream of the orifice and continues to increase in size until the air phase begins to protrude inside the orifice to form its own bulb. The air phase’s bulb begins to choke the orifice causing the collagen bulb to form a thin neck (inside orifice). As the air phase bulb grows in size, it eventually occupies the entire orifice which then perturbs/dissects the elongated collagen neck resulting in a collagen droplet. The pressure or flow rate at which the air phase is supplied can be changed to affect the size of the produced bubbles which in turn provides a facile method to control the size of the collagen droplets (FIG. 5C). Particle sizes between about 50 to about 300 microns have been studied but it will be understood the device geometry can be readily modified for production of smaller and larger particles outside the above range. According to a nonlimiting embodiment, the collagen flow rate may be between the range of about 0.7 to 2.4 ml/hr, the oil flow rate may be between the range of about 1 .8 to 6 ml/hr and the air pressure may be between the range of about 2 to 5 PSI. For example, by varying the pressure and flow rate, the air bubbles may have a size ranging 50 pm to 300 pm resulting into collagen droplets having a size ranging between the range of 50 pm to 300 pm.

The continuous oil flow carrying the collagen solution droplets and air bubbles produced by each flow focusing units enters a common collection microchannel via through-holes. An external tube connected to the collection microchannel allows retrieval of the prepared emulsion and is connected to another flowing oil (basic oil) phase downstream using a T-connector (FIG. 5D). This basic oil contains an oil dissolvable base (0.3% vol/vol triethylamine, [pH 11]). Triethylamine molecules diffuse into the collagen droplets and increase the pH of the initially acidic collagen solution to a neutral, inducing fibrillogenesis and gelation into collagen microgels. The vol% of the oil dissolvable base allows adjusting the final pH of the collagen solution which can be leveraged to attain different collagen fiber sizes. The emerging collagen microgels are flowing inside the tube at 37°C for about 1 hour to expedite fibrillogenesis.

The produced droplets are washed from the continuous oil phase using five rinsing cycles with tween solution (0.1 % vol/vol in DI water). The microgels and tween solution are first mixed with a vortex machine and then later centrifuged (2000 Gs) to separate the microgels. Each time the supernatant is aspirated, and a fresh solution of Tween is added. Once all the oil is removed with tween washes, the microgels are washed another five times in DI water to remove the tween. The washed microgels are then stored in DI water at 4°C for later use. These washed droplets are monodisperse and have a narrow size distribution with coefficient of variation of 5.5% (FIG. 5E).

The microparticles produced using the mentioned technique can then be ice templated (frozen) at different temperatures to attain varying external/internal porosity values (FIG. 6A). The initial collagen concentration affects the final fiber density as well as the particle morphology resulting after ice templating. Furthermore, for certain conditions different pore sizes can result in the core and the periphery the particle. For example, for the freezing rate of 10°C/min, the particle develops a thin shell which separates the smaller ice crystal templated core from the larger ice crystal templated exterior. These particles can be lyophilized immediately after the freezing step/ice templating to preserve them for later use.

Transmission electron micrographs reveal that the collagen particles produced using the above-mentioned microfluidic technique contain selfassembled collagen fibers with diameters ranging from about 30nm to about 500nm, wherein the fibers exhibit a periodic D-banding structure with about 67nm wavelength (FIG. 6B). Furthermore, significant collagen compaction takes places during ice templating which leads to the formation of thin internal walls. The thickness of these walls decreases with increasing freezing rate.

Various methods of crosslinking may be used including chemically crosslinking, physical cross linking or enzymatic crosslinking the collagen molecules. Chemical crosslinking may be performed using, but not limited to, genipin, 1-Ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC), or dialdehyde starch (DAS). Physical crosslinking may be done by UV crosslinking using riboflavin. Enzymatic crosslinking may be done using microbial transglutaminase (mTG) or Lysyl oxidase (LOX) to give some non-limiting examples, see for example K. Adamiak, A. Sionkowska I International Journal of Biological Macromolecules, 161 , (2020) 550-560.

As shown in FIG. 6C, the stiffness or Young’s modulus of the particle can be measured using atomic force microscopy. Furthermore, by employing crosslinking techniques such as dehydrothermal treatment (DHT) at 120°C for 24 hours in 50 mTorr vacuum and 0.5% vol/vol glutaraldehyde (GTA), a wide range of particle stiffnesses can be achieved. For crosslinking with GTA, the collagen particles are mixed with the crosslinker solution before the ice templating step which improves structure retention post thawing.

Using custom-made poly(dimethylsiloxane) (PDMS) molds, square constructs of 10mm x 10mm with 1 mm thickness were casted. These constructs were made using cellular non templated particles, non templated particles crosslinked with GTA and ice templated (1 °C/min) particles supplemented with a 10mg/mL fibrinogen solution, and were crosslinked with a 500 U/mL thrombin solution. The ucMSCs were included in all three cases at a final density of 1 xio ⁶ cells/mL. The three samples were cultured for >2 weeks, and images were recorded every 24 hours to compare the contraction of the square constructs over the culture period.

The sample images of the construct at days 1 , 5, 10 and 15 (FIG. 7A). Between the three microparticle-based samples, the macroscopic images indicated the non templated crosslinked with GTA and 1 °C/min collagen particles sample to contract insignificantly, whereas the non templated sample started shown contraction by day 10. The difference in the improved mechanical properties is due to the GTA crosslinking and high compaction of collagen in the non templated and ice templated particles, respectively.

To establish cytocompatibility of these particles, umbilical cord derived mesenchymal stromal/stem cells (ucMSC) were stained with a fluorescent LIVE/DEAD kit. The viability analysis and a representative micrograph of ucMSC on collagen-based microparticles are shown in FIG. 7B. In all cases the cells exhibited >80% viability after 3 days of culture.

The cell morphology for the three cases were also studied at day 21 using fluorescent immunostaining, where the cells were fixed and the F-actin filaments of their cytoskeleton, nuclei stained with Hoechst, smooth muscle actin filaments and Type 1 collagen as illustrated in FIG. 7C. It can be observed that all particles allow long term culture and proliferation of cells. Higher magnification images (right) reveal that the overall morphology of the particles is maintained at day 21 . References

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