EP4164706A1 - Vorrichtung und verfahren zur gewebezüchtung und regenerativen medizin - Google Patents

Vorrichtung und verfahren zur gewebezüchtung und regenerativen medizin

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Publication number
EP4164706A1
EP4164706A1 EP20816241.2A EP20816241A EP4164706A1 EP 4164706 A1 EP4164706 A1 EP 4164706A1 EP 20816241 A EP20816241 A EP 20816241A EP 4164706 A1 EP4164706 A1 EP 4164706A1
Authority
EP
European Patent Office
Prior art keywords
preferentially
cell
microcarrier
cells
culture
Prior art date
Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
Pending
Application number
EP20816241.2A
Other languages
English (en)
French (fr)
Inventor
Daniel Naveed TAVAKOL
Josefine TRATWAL
Olaia Naveiras
Thomas Braschler
Fabien BONINI
Amélie Barbara H. BÉDUER
Martina GENTA
Joé BREFIE-GUTH
Current Assignee (The listed assignees may be inaccurate. Google has not performed a legal analysis and makes no representation or warranty as to the accuracy of the list.)
Ecole Polytechnique Federale de Lausanne EPFL
Universite de Geneve
Universite de Lausanne
Centre Hospitalier Universitaire Vaudois CHUV
Original Assignee
Ecole Polytechnique Federale de Lausanne EPFL
Universite de Geneve
Universite de Lausanne
Centre Hospitalier Universitaire Vaudois CHUV
Priority date (The priority date is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the date listed.)
Filing date
Publication date
Application filed by Ecole Polytechnique Federale de Lausanne EPFL, Universite de Geneve, Universite de Lausanne, Centre Hospitalier Universitaire Vaudois CHUV filed Critical Ecole Polytechnique Federale de Lausanne EPFL
Publication of EP4164706A1 publication Critical patent/EP4164706A1/de
Pending legal-status Critical Current

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    • C—CHEMISTRY; METALLURGY
    • C12—BIOCHEMISTRY; BEER; SPIRITS; WINE; VINEGAR; MICROBIOLOGY; ENZYMOLOGY; MUTATION OR GENETIC ENGINEERING
    • C12N—MICROORGANISMS OR ENZYMES; COMPOSITIONS THEREOF; PROPAGATING, PRESERVING, OR MAINTAINING MICROORGANISMS; MUTATION OR GENETIC ENGINEERING; CULTURE MEDIA
    • C12N5/00—Undifferentiated human, animal or plant cells, e.g. cell lines; Tissues; Cultivation or maintenance thereof; Culture media therefor
    • C12N5/0068—General culture methods using substrates
    • C12N5/0075—General culture methods using substrates using microcarriers
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L27/00—Materials for grafts or prostheses or for coating grafts or prostheses
    • A61L27/50—Materials characterised by their function or physical properties, e.g. injectable or lubricating compositions, shape-memory materials, surface modified materials
    • A61L27/52—Hydrogels or hydrocolloids
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L27/00—Materials for grafts or prostheses or for coating grafts or prostheses
    • A61L27/50—Materials characterised by their function or physical properties, e.g. injectable or lubricating compositions, shape-memory materials, surface modified materials
    • A61L27/54—Biologically active materials, e.g. therapeutic substances
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L27/00—Materials for grafts or prostheses or for coating grafts or prostheses
    • A61L27/50—Materials characterised by their function or physical properties, e.g. injectable or lubricating compositions, shape-memory materials, surface modified materials
    • A61L27/56—Porous materials, e.g. foams or sponges
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L2300/00—Biologically active materials used in bandages, wound dressings, absorbent pads or medical devices
    • A61L2300/40—Biologically active materials used in bandages, wound dressings, absorbent pads or medical devices characterised by a specific therapeutic activity or mode of action
    • A61L2300/412—Tissue-regenerating or healing or proliferative agents
    • A61L2300/414—Growth factors
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L2300/00—Biologically active materials used in bandages, wound dressings, absorbent pads or medical devices
    • A61L2300/40—Biologically active materials used in bandages, wound dressings, absorbent pads or medical devices characterised by a specific therapeutic activity or mode of action
    • A61L2300/432—Inhibitors, antagonists
    • C—CHEMISTRY; METALLURGY
    • C12—BIOCHEMISTRY; BEER; SPIRITS; WINE; VINEGAR; MICROBIOLOGY; ENZYMOLOGY; MUTATION OR GENETIC ENGINEERING
    • C12N—MICROORGANISMS OR ENZYMES; COMPOSITIONS THEREOF; PROPAGATING, PRESERVING, OR MAINTAINING MICROORGANISMS; MUTATION OR GENETIC ENGINEERING; CULTURE MEDIA
    • C12N2533/00—Supports or coatings for cell culture, characterised by material
    • C12N2533/30—Synthetic polymers
    • C12N2533/40—Polyhydroxyacids, e.g. polymers of glycolic or lactic acid (PGA, PLA, PLGA); Bioresorbable polymers
    • C—CHEMISTRY; METALLURGY
    • C12—BIOCHEMISTRY; BEER; SPIRITS; WINE; VINEGAR; MICROBIOLOGY; ENZYMOLOGY; MUTATION OR GENETIC ENGINEERING
    • C12N—MICROORGANISMS OR ENZYMES; COMPOSITIONS THEREOF; PROPAGATING, PRESERVING, OR MAINTAINING MICROORGANISMS; MUTATION OR GENETIC ENGINEERING; CULTURE MEDIA
    • C12N2533/00—Supports or coatings for cell culture, characterised by material
    • C12N2533/50—Proteins
    • C12N2533/54—Collagen; Gelatin
    • C—CHEMISTRY; METALLURGY
    • C12—BIOCHEMISTRY; BEER; SPIRITS; WINE; VINEGAR; MICROBIOLOGY; ENZYMOLOGY; MUTATION OR GENETIC ENGINEERING
    • C12N—MICROORGANISMS OR ENZYMES; COMPOSITIONS THEREOF; PROPAGATING, PRESERVING, OR MAINTAINING MICROORGANISMS; MUTATION OR GENETIC ENGINEERING; CULTURE MEDIA
    • C12N2533/00—Supports or coatings for cell culture, characterised by material
    • C12N2533/70—Polysaccharides

Definitions

  • This invention generally pertains to the tissue engineering and regenerative medicine fields.
  • Biomaterials are increasingly used for in-vivo cell delivery (some examples: US2017136153A1, W02012149358 Al, W02013036875A1), where they often take the name of scaffolds(Garg et al., 2012). Besides offering adhesion sites for anchorage- dependent cells, they allow additional functionalities such as defined co-culture or pre vascularization (W02008008229A2). An interesting alternative is also to deliver cells and scaffolds separately. This is possible for mobile cells such as hematopoietic stem cells, when strong homing factors are known (W02017/136837A1). Scaffolds can further be composed of a variety of synthetic or natural polymers, but also decellularized extracellular matrix (for example, W02017/223529, W02016168752 Al).
  • Injectable scaffolds for example preformed large scaffolds(Bencherif et al., 2012a), W02012149358 Al, but also other injectables, such as viscoelastic hydrogels (Nih et al., 2017) or cements (for example W02014160232A2) are of particular utility since they permit minimally invasive delivery.
  • injectables such as viscoelastic hydrogels (Nih et al., 2017) or cements (for example W02014160232A2) are of particular utility since they permit minimally invasive delivery.
  • For scaffold injection it is sometime necessary to dehydrate scaffolds; this is typically obtained by applying some sort of mechanical pressure during injection (Bencherif et al., 2012a) and/or partial dehydration before actual injection (Beduer et al., 2015a).
  • Microcarriers are particles on the scale of tens of micrometers to the millimeter range. They are intended for the culture of adherent cells in suspension. Many different microcarriers are commercially available (the Cytodex, Cytopore and Cultisphere series being among the more renowned), and many fabrication methods and compositions established in the scientific and patent literature (example: EP1801122A1; defined shapes in W02014037862A1). Microcarrier-based cell culture is also a well-established art.
  • microcarriers are large-scale production of cells or other biological such as antibodies or cells (example: hematopoietic stem cells, W00046349A1 and W02012127320A1); this can elegantly be combined with cell harvest by affinity to microcarriers (W02009/002456 A2).
  • Microcarriers are also increasingly used to deliver cell therapy(example: US2015361395A1), since they can provide cells with adhesion during the transplantation process (Newland et al., 2015).
  • the necessity for a defined transition with substantial volume reduction from a dilute suspension to a dense, transplantable injectable is particularly a problem for cells expanded to high numbers (typically, various stem cells such as embryonic stem cells, induced pluripotent stem cells, or endogenous stem cells such as hematopoietic stem cells and lymphoid progeny or neural stem cells), but also for postmitotic cells with a high-demand metabolism (for example, transient amplifying or progenitor populations, neurons, transport epithelia, cardiomyocytes or immune cells).
  • various stem cells such as embryonic stem cells, induced pluripotent stem cells, or endogenous stem cells such as hematopoietic stem cells and lymphoid progeny or neural stem cells
  • postmitotic cells with a high-demand metabolism for example, transient amplifying or progenitor populations, neurons, transport epithelia, cardiomyocytes or immune cells.
  • microcarriers are cultured in large, dilute volumes to satisfy high metabolic needs arising from stem and progenitor cells, neurons, cardiomyocytes, transport epithelial cells or other highly active cells, and where in a second step injectable implants with high mechanical stability arising from partial microcarrier compression need to be produced.
  • the application is to simultaneously transplant the cells and form mechanically stable local niche.
  • the aspiration pressure applied to the interstitial fluid determines the final microcarrier density and indirectly the mechanical properties of the final implant.
  • the dehydration rate as expressed by the fluid flow rate through relevant portions or all of the microcarrier material being compacted. As cells are sensitive to shear(Tanzeglock et al., 2009), this must be controlled if cell presence, integrity and if applicable cell interrelations are to be preserved.
  • the compacted microcarriers need to remain with either directly a suitable syringe or a transfer device, rather than sticking to the fluid removal device, since otherwise it needs to be scraped of a membrane with concomitant risk of loss of cell viability, and generation of foreign particles.
  • the present invention addresses these challenges.
  • microcarriers lend themselves to the present invention, there are differences in their properties that make their ease of use or performance different. Scaffolds with hard, spherical particles yield relatively brittle pastes that are nearly unable to sustain shapes other than externally imposed ones; such an example is provided in the literature by partial rehydration for culture of metabolically relatively inactive cells (Fig. 8h in (Xia et al., 2017)), but can also be reached from suspension culture with the present invention by using for example cytodex microcarriers. In this case, the primary advantage of dehydration lies in the cell density.
  • Softer, porous materials for example the commercial material cultisphere G or possibly newer non-porous (EP1801122A1) or porous gelatin carriers (EP3210634A1) make for more stable biomaterial pastes, most likely due to better shape-adaption and particle interlocking.
  • the invention is particularly useful in conjunction with microcarriers exhibiting a plateau in compression such as the ones describes in W02017/029633; these are able to maintain their mechanical properties even for major volume changes; once adjusted these are resilient to loss of pore fluid by evaporation or dilution by addition of small to moderate amounts of liquid, for instance by contact with body fluids.
  • Scaffolds have also been used to provide stromal functionality, for example by allowing implantation of mesenchymal stem cells for nerve repair (CN104491925B), or for co culture and co-transplantation of stem cells with stromal components, for instance to mimick a bone marrow environment for acute myeloid leukemia cancer stem cells, hematopoietic stem cells or other stem cells along with a mesenchymal stem cells cultured on the scaffold to provide stromal support (US2011/0207166A1).
  • the present invention can also be advantageously used for such applications, provided the co-culture can be implemented on microcarriers suitable for this invention.
  • a possible application of the present invention is the generation of an ectopic or new marrow niche for hematopoietic stem cells by means of a co-transplantation with stromal cells on microcarriers. It is indeed the basis of hematopoietic stem cell transplantation that hematopoietic stem cells are extraordinarly dependent on their niche for proliferation and maintenance. In fact, to guarantee successful engraftment, at least partial emptying of the hematopoietic niche of the recipient is at present necessary. Chemotherapy and whole-body irradiation are the major means to this end, although niche manipulation by injection of supplementary mesenchymal stem cells (W02016151476A1) together with the use of mobilizing agents such as GM-CSF are also being explored.
  • An alternative means is to provide a local, possibly ectopic niche altogether (W02017136837A1, filled by separate transplantation of hematopoietic stem cells).
  • the teachings of this invention can be used to create an injectable, yet coherent and shape-stable implant from a microcarrier-based co-culture of stromal cells (such as the OP9 line) and hematopoietic stem/progenitor cells after efficient in-vitro expansion under microcarrier culture conditions (cytokine-based, see for example W02017075389A1, or based solely on the action of the OP-9 or other stromal cells).
  • the present inventors developed a process and device for direct in- vitro to in-vivo transfer of a living microcarrier culture by controlled dehydration into a viscoelastic material capable of cell-protection during delivery.
  • One particular purpose of the present invention is that of providing a process and device for direct in-vitro to in- vivo transfer of living microcarrier co-cultures.
  • the invention as described in here serves to compact a microcarrier culture or co-culture into an injectable implant with a paste-like consistency.
  • the final implant remains injectable, but due to partial dehydration, has some mechanical strength, as evident for instance by rheometric G’ values in the range of lOPa to IMPa, more preferentially lOOPa to lOkP and most preferentially 500Pa to 5kPa.
  • the partial dehydration system consists of a fluid drain capable of absorbing amounts at least equal, but typically by far (lOx or more) exceeding the amount of fluid to be absorbed from the microcarrier culture. It further is capable of separately, and stringently, regulate both final interstitial fluid pressure to be achieved (between 20 Pa and lOkPa, preferentially 50 Pa to lkPa, and even more preferentially between 100 Pa and 800 Pa below atmospheric pressure), and the maximum fluid flow rate experienced by microcarriers to subsequently be delivered in-vivo (or for in-vitro applications such as 3D printing) in the range between O.Olmm/s and lOcm/s, more preferentially between O.lmm/s and 5cm/s, and even more preferentially between 0.5mm/s and lOmm/s.
  • the invention further comprises a transfer recipient.
  • the dehydration process is initiated when the transfer device is loaded with the microcarrier culture, optionally through a filling column.
  • the flow rate can be imposed by the dehydration system itself, or optionally, by a self-assembling dense plug of microcarrier material at the orifice of the transfer device in contact with the dehydration device. In this case, the dimensions of the orifice are key for the dehydration process.
  • the desired final pressure which can optionally be monitored, for example by a pressure gauge or glass capillary in touch with the material being dehydrated
  • the compacted, paste-like microcarrier culture is confined in the transfer device and can be manipulated or transported.
  • the transfer device typically ensures connection to an injection syringe, or is itself an adapted syringe (with an air vent for piston insertion), and also ensures connection to the delivery tubing (typically, a catheter or blunt needle), through which the compacted, living microcarrier culture or co-culture is then delivered in-vivo. It is optionally possible to sub-sample some the compacted material to ensure appropriate composition and viability prior to injection.
  • a “scaffold” is any three dimensional material having a framework architecture, for instance a support structure comprising hollow spaces within it.
  • a scaffold is an artificial structure capable of supporting cell culture and/or three- dimensional tissue/organ formation in vivo, in vitro or ex vivo.
  • a scaffold is also referred herewith as a “biomaterial” or “bioscaffold”.
  • a scaffold can be considered the physical structure (including biodegradable and/or permanent materials) upon which or into which cells directly or indirectly associate or attach.
  • a bioscaffold may allow or facilitate cell attachment and migration, delivers and retains cells and biochemical factors, enables diffusion of vital cell nutrients and expressed products, exerts certain mechanical and biological influences to modify the behaviour of the cell phase and so forth.
  • a “polymeric material” is any material comprising polymers, large molecules (also known as macromolecules) composed of many repeated smaller units, or subunits, called monomers, tightly bonded together by covalent bonds.
  • gel refers to a non-fluid colloidal network or polymer network that is expanded throughout its whole volume by a fluid.
  • a gel is a solid three-dimensional network that spans the volume of a liquid medium and ensnares it through surface tension effects.
  • the internal network structure may result from physical bonds (physical gels) or chemical bonds (chemical gels) such as covalent, ionic, hydrogen and/or Van der Waals bonds.
  • a “microcarrier” is a particle on the scale of few micrometers up to millimeters, intended for the culture of adherent cells in suspension.
  • Microcarriers can comprise or be substantially composed of polymeric materials such as plastic materials, biopolymers and/or ceramic materials.
  • Many different microcarriers are commercially available (the Cytodex, Cytopore and Cultisphere series being among the more renowned), and many fabrication methods and compositions established in the scientific and patent literature (example: EP1801122A1). Microcarrier-based cell culture is also a well-established art.
  • microcarriers are large-scale production of cells or other biological such as antibodies (example: hematopoietic stem cells, W00046349A1); this can elegantly be combined with cell harvest by affinity to microcarriers (W02009/002456 A2).
  • Microcarriers are also increasingly used to deliver cell therapy (example: US2015361395A1), since they can provide cells with adhesion during the transplantation process (Newland et al., 2015).
  • microcarriers according to the invention are substantially composed out of hydrogel.
  • hydrogel refers to a gel in which the swelling agent is water.
  • a hydrogel is a macromolecular polymer gel constructed of a network of cross-linked polymer chains. It is synthesized from hydrophilic monomers, sometimes found as a colloidal gel in which water is the dispersion medium.
  • Hydrogels are highly absorbent (they can contain up to over 90% water) natural or synthetic polymeric networks. As a result of their characteristics, hydrogels develop typical firm yet elastic mechanical properties. Hydrogels have been used in biomedical applications, such as contact lenses and wound dressings.
  • hydrogels are more biocompatible than hydrophobic elastomers and metals. This biocompatibility is largely due to the unique characteristics of hydrogels in that they are soft and contain water like the surrounding tissues and have relatively low frictional coefficients with respect to the surrounding tissues. Furthermore, hydrogels permit diffusion of aqueous compositions, and the solutes, there through, and have a high permeability to water and water- soluble substances, such as nutrients, metabolites and the like.
  • hydrogel pore radius, morphology, or its permeability may change the hydrogel pore radius, morphology, or its permeability to different molecular weight proteins.
  • volume or dimensions (length, width, and thickness) of a hydrogel can be selected based on the user’s needs, such as e.g. the region or environment into which the hydrogel is to be implanted in the frame of a surgical setting.
  • the mechanical properties of the material can be tailored according to the application site by changing the hydrogel composition (molecular chain length, crosslinking, water content and the like).
  • suitable materials constituting the microcarriers to be used in the frame of this invention include natural polymers, such as polysaccharides, co polymers of polysaccharides (cellulose, agarose, alginate, starch, chitosan and many others), polypeptides (silk, collagen, gelatin and many others), amelogenin or synthetic polymers such as polyurethanes, poly-olefins, polyethylene glycol (PEG), poly(glycolide) (PGA), poly(L-lactide) (PLA), carboxymethylcellulose (CMC) or poly(lactide-co-glycolide) (PLGA).
  • natural polymers such as polysaccharides, co polymers of polysaccharides (cellulose, agarose, alginate, starch, chitosan and many others), polypeptides (silk, collagen, gelatin and many others), amelogenin or synthetic polymers such as polyurethanes, poly-olefins, polyethylene glycol (P
  • the microcarrier may also comprise either at least one glycosaminoglycane or at least one proteoglycane, or a mixture of those two substances.
  • the glycosaminoglycane may be for example a hyaluronic acid, chondroitinsulfate, dermatansulfate, heparansulfate, heparine or keratansulfate.
  • the ability to use different materials is useful in different applications and adds a further degree of versatility to the device and methods described herein.
  • the base material is not limiting as long as the other essential mechanical requirements are met, and the microcarriers can withstand a (partial) dehydration process.
  • the degradation/resorption rate of the carrier upon in vivo application/implant in a host is mainly dependent on physico-chemical properties of the polymeric material of which it is composed of, as well as further factors such as crosslinking of the polymers, the polymer concentration, the site of implant into a host and the like.
  • the degradation/resorption rate can be calibrated by adjusting said physico chemical parameters, such as for instance by polymer crosslinking (if present), the use of inhibitor molecules, by changing the polymer density, crystallinity and/or its molecular weight distribution, changing the materials’ porosity and so forth.
  • the scaffold may be, at least in part and at least in some portion thereof, intrinsically biodegradable in vivo.
  • Microcarriers can exist in different shapes and porosities.
  • microcarriers according to the invention are highly porous and with an irregular shape.
  • a preferred microcarrier in the frame of the invention comprises pores that are interconnected in order to create a continuous net of material that can act as a plausible physical support for elements such as cells or bioactive agents, while providing at the same time additional key features to the scaffold such as its softness, low resistance to interstitial flow, high compressibility, outstanding ability to regulate the capillary pressure to a constant level over a wide range of hydration states, easiness of cell/tissue invasion and so forth.
  • the porosity of the material is preferably comprised between 50% and 99%, allowing for evacuation of liquid from the pores upon compression.
  • Non-porous and/or regular shape microcarriers may be envisaged for use according to the invention.
  • the pore size is typically comprised between lpm and 10mm, preferably between lpm and 5mm, more preferably between lpm and 2mm, even more preferably between 5pm and 500pm, obviously the size of the microcarrier itself being the upper limit of a pores size.
  • This range of pore size is particularly convenient for a scaffold material for use in tissue engineering or regenerative medicine, since it is e.g. high enough to enable the growth of vessels through the porous material.
  • the overall polymer content is comprised between 0.01% and 100% in mass of dry polymer material, preferably between 0.5% and 3%.
  • liquid is removed therefrom and the hydration level decreases.
  • Microcarriers are typically used in suspension state, wherein two main regimens exist: in a dilute regime, microcarriers are separated from each others, whereas in a dense regime, a contiguous network of neighboring microcarriers is formed.
  • the progressive transition from a dilute regime to a dense regime is called “jamming transition”, and the microcarrier concentration associated to said jamming transition can be approximately found, as will be apparent to a person skilled in the art performing a sedimentation assay.
  • the concentration at the jamming transition state can be considered as the full hydration state of a plurality of microcarriers (100% hydration).
  • a “soft” material is any material that is either compressible, reversibly compressible, flexible, elastic or any combination thereof.
  • the scaffold In order to be used in living subjects, moreover, the scaffold must also imperatively be a biocompatible and/or sterilisable material suitable for medical uses.
  • a microcarrier and/or suspension according to the invention typically behaves, depending on the hydration state, as a viscoelastic material, particularly in the dense state of suspension as defined above (100% hydration or less). Depending on the shape, porosity, density and intrinsic material properties, an elastic response can be observed for low deformations whereas a viscous behavior is observed upon high deformations.
  • the elastic regime is important for shape stability, whereas the viscous regime is considered important for injectability and shapeability of an implant.
  • the scaffold material can be functionalized with additional elements such as for instance bioactive molecules.
  • Said elements can be coated on or embedded within the obtained scaffold with any suitable means known in the art, and can provide additional functional properties to the material such as enhanced/reduced biodegradation, physical stabilization, biological activity and the like.
  • a “bioactive molecule”, as well as “(bio)active compound” or “therapeutic agent”, is any active agent having an effect upon a living organism, tissue, or cell. The expression is used herein to refer to any compound or entity that alters, inhibits, activates, or otherwise affects biological or chemical events.
  • bioactive compounds include, but are not limited to, a small molecule, a growth factor, a protein, a peptide, an enzyme, an antibody or any derivative thereof (such as e.g.
  • multivalent antibodies multispecific antibodies, scFvs, bivalent or trivalent scFvs, triabodies, minibodies, nanobodies, diabodies etc.
  • an antigen e.g., a nucleic acid sequence (e.g., DNA or RNA), a hormone, an anti inflammatory agent, an anti-viral agent, an anti-bacterial agent, a cytokine, an oncogene, a tumor suppressor, a transmembrane receptor, a protein receptor, a serum protein, an adhesion molecule, a lypidic molecule, a neurotransmitter, a morphogenetic protein, a differentiation factor, an analgesic, organic molecules, metal particles, radioactive agents, polysaccharides, a matrix protein, and any functional fragment or derivative of the above, as well as any combinations thereof.
  • a functional fragment is herein meant any portion of an active agent able to exert its physiological/pharmacological activity.
  • a functional fragment of an antibody could be an Fc region, an Fv region, a Fab/F(ab’)/F(ab’)2 region and so forth.
  • the bioactive compounds can be added to the microcarrier material using any suitable method known in the art, such as surface absorption, physical immobilization, e.g., using a phase change to entrap the substance in the scaffold, and the like.
  • a growth factor can be mixed with the microcarrier material composition.
  • covalent coupling e.g., using alkylating or acylating agents, is used to provide a stable, long-term presentation of a bioactive substance on the microcarrier in a defined conformation.
  • non-covalent adsorbtion can be used, for example electrostatic, hydrophobic, dipole-dipole, hydrogen bonding and the like.
  • the microcarrier are basically cell-free, but they are later on seeded in vitro or in vivo with cells after their production with the intention to transplant cells in a tissue engineering/regenerative medicine approach.
  • Cell seeding in its simplest implementation, involves placing the microcarriers in a cell suspension and letting the cells adhere and colonize the microcarrier.
  • compressible microcarriers it is possible to greatly enhance the cell seeding process by partially dehydrating (compressing) the scaffold prior to addition of the cell suspension. In this case, the aspiration force created by the compressed microcarries tendency to expand draws the cell suspension into the microcarrier, and will seed cells deeply into the microcarrier.
  • an external perfusion device to induce flow across the microcarrier and use this flow to efficiently seed cells.
  • cell-adhesivity and/or porosity of the microcarrier plays a crucial role in organizing the cell seeding.
  • Cells of a given type will only adhere in areas which are reachable for them by fluid flow, or, on a longer term, migration, and which allow for their attachment.
  • a porous microcarrier it is possible to sterically entrap cells even when cells do not adhere to the microcarrier material per se.
  • one cell type can mediate the attachment of another cell type which is not normally capable of adhesion to the microcarrier.
  • Areas of very low pore size will act as a complete barrier for cell seeding by flow, be it by external perfusion or by intrinsically generated flow due to microcarrier expansion. This is the case of non-porous microcarriers, where only the microcarrier surface would be colonized by seeded cells.
  • Different cell types or tissues fragments e.g., stem vs. differentiated, and/or with various phenotypes in terms of differentiation, activation, metabolic or functional state, are optionally co-resident in the microcarrier of the invention.
  • the microcarriers should be chosen to be suitable for use with any cell type or tissue fragments that one may want to transplant.
  • Such cells include but are not limited to, various stem cell populations (embryonic stem cells differentiated into various cell types), bone marrow or adipose tissue derived adult stem cells, mesenchymal stem cells, cardiac stem cells, pancreatic stem cells, neuronal cells, glial cells, spermatozoids and ovocytes, endothelial progenitor cells, outgrowth endothelial cells, dendritic cells, hematopoietic stem cells, neural stem cells, satellite cells, side population cells.
  • Such cells may further include but are not limited to, differentiated cell populations including osteoprogenitors and osteoblasts, chondrocytes, keratinocytes for skin, intestinal epithelial cells, smooth muscle cells, cardiac muscle cells, epithelial cells, endothelial cells, urothelial cells, fibroblasts, myoblasts, chondroblasts, osteoclasts, hepatocytes, bile duct cells, pancreatic islet cells, thyroid, parathyroid, adrenal, hypothalamic, pituitary, ovarian, testicular, salivary gland cells, adipocytes and combinations thereof.
  • differentiated cell populations including osteoprogenitors and osteoblasts, chondrocytes, keratinocytes for skin, intestinal epithelial cells, smooth muscle cells, cardiac muscle cells, epithelial cells, endothelial cells, urothelial cells, fibroblasts, myoblasts, chondroblasts, osteoclasts, hepatocytes,
  • smooth muscle cells and endothelial cells may be employed for muscular or tubular scaffolds, e.g., scaffolds intended as vascular, esophageal, intestinal, rectal, or ureteral scaffolds; chondrocytes may be employed in cartilaginous scaffolds; cardiac muscle cells may be employed in heart scaffolds; hepatocytes and bile duct cells may be employed in liver scaffolds; myoblasts may be used in muscle regeneration; epithelial, endothelial, fibroblast, and nerve cells may be employed in scaffolds intended to function as replacements or enhancements for any of the wide variety of tissue types that contain these cells.
  • scaffolds intended as vascular, esophageal, intestinal, rectal, or ureteral scaffolds e.g., chondrocytes may be employed in cartilaginous scaffolds
  • cardiac muscle cells may be employed in heart scaffolds
  • hepatocytes and bile duct cells may be employed in liver scaffolds
  • myoblasts may be used in muscle
  • microcarriers of the invention may comprise any cell population competent to participate in regeneration, replacement or repair of a target tissue or organ, particularly barely-reachable ones such as bone marrow, kidneys, brain, lungs or pancreas.
  • a target tissue or organ particularly barely-reachable ones such as bone marrow, kidneys, brain, lungs or pancreas.
  • Microcarriers according to the invention may be prepared or chosen so that the degradation time may be controlled by using a mixture of degradable components in proportions to achieve a desired degradation rate.
  • the cells themselves aid in the degradation.
  • scaffold compositions are sensitive to degradation by materials secreted by the cells themselves that are seeded within the scaffold.
  • microcarriers according to the invention can be ideally transplanted on or close to a target tissue, introduced into or onto a bodily tissue/organ using a variety of methods and tools known in the art, preferably via minimally invasive surgical devices and procedures that envisage an injection step by using cannulas, syringe needles or catheters, preferably blunt ended catheters.
  • the terms "patient” or “subject” refer to an animal, preferably to a mammal, even more preferably to a human, including adult and child. However, the terms can also refer to non-human animals, in particular mammals such as dogs, cats, horses, cows, pigs, sheeps, rodents and non-human primates, among others, that are in need of treatment. In some instances, a “subject” can be a plant.
  • treatment generally refers to any act intended to ameliorate the health status of subjects such as an animal, particularly a mammal, more particularly a human, such as therapy, prevention, prophylaxis and retardation of the disease, and includes: (a) inhibiting the disease, i.e., arresting its development; or (b) relieving the disease, i.e., causing regression of the disease and/or its symptoms or conditions such as improvement or remediation of damage.
  • such term refers to the amelioration or eradication of a disease or symptoms associated with a disease.
  • this term refers to minimizing the spread or worsening of the disease resulting from the administration of one or more therapeutic agents to a subject with such a disease.
  • prevention or “preventing” relates to hampering, blocking or avoid a disease from occurring in a subject which may be, for any reason, predisposed to the disease but has not yet been diagnosed as having it for example based on familial history, health status or age.
  • dehydration device allowing to apply controlled pressure and flow rate during the dehydration procedure, with a transfer device that can, but does not necessarily have to be a syringe like device.
  • the dehydration device or in the following also dehydration “box”, must be able to control final pressure and flow rate during the dehydration.
  • an elementary implementation with a capillary conductor and a reservoir is given, but other solutions are also possible, for example with flow regulators driven by pressure regulators in an automated feedback loop.
  • a separate transfer device for transfer of the compacted microcarrier culture to an injection device such as a syringe.
  • the transfer device may however also be directly integrated with the syringe or other in-vivo delivery device, and if cell and material losses and possible foreign particle contamination can be accepted, manual transfer would also be possible.
  • Example 1 Implementation of a dehydration device, transfer device and usage for injection
  • This example provides a basic implementation of the dehydration device, transfer tip and filling column. It also describes the dehydration process used. It makes use a specific dehydration device, which has been designed to dehydrate a solution of microcarriers at an optimal interstitial pressure for injection (and thus optimal polymer concentration). It allows an efficient removal of the dead volume while guaranteeing a safe, reproducible and easy transition from the in vitro cell culture to the in vivo injection. The whole process is harmless for the cells to be transplanted. Dehydration is complete in few minutes and the compacted microcarriers are then ready to be injected in vivo. The whole system is autoclavable (all pieces are made from polypropylene plastic) and can be treated for endotoxin removal process (with concentrated sodium hydroxide) in order to prevent excessive inflammatory response from the host.
  • endotoxin removal process with concentrated sodium hydroxide
  • Fig. 1 shows the operating principle of the drying device.
  • the sample is applied in a drying column.
  • a pre-set aspiration pressure is generated by a lower level of fluid in the waste reservoir of the drying box, and is transmitted to the sample via a capillary conductor (cloth).
  • the magnitude of the aspiration pressure is set by the level of fluid in the waste reservoir, and changes little during the dehydration process since the reservoir capacity by far exceeds the liquid volume to be removed.
  • the direction of the fluid movement is determined by the relative magnitude of the elastic expansion force of the microcarrier sample in the drying column as compared to the imposed aspiration pressure. Therefore, after equilibration, the aspiration pressure determines the fluid content of the microcarrier sample regardless of the initial fluid content.
  • the fluid flow rate is determined by the difference of expansion pressure and hydrostatic under-pressure, but also by the total flow resistance.
  • the primary source of flow resistance is the microcarrier material near the tip of the transfer column.
  • Fig. 2 provides an overview over the process flow used to obtain an injectable from microcarrier cultures.
  • the starting material in this embodiment is a dilute suspension of carboxymethylcellulose microcarriers (CCM, synthesis given in Example 2) colonized with the cells of interest (typically, but by no means necessarily, a co-culture of HSPC and OP-9 feeder cells, Fig. 2A), but the principle is the same with any microcarrier culture (cells and carriers).
  • CCM carboxymethylcellulose microcarriers
  • Fig. 2A a co-culture of HSPC and OP-9 feeder cells
  • This dilute suspension is transferred to the dehydration column, which consists itself of two parts, namely the loading column and the transfer tip (Fig. 2B). Excess fluid is drained into the waste reservoir by the dehydration box, until equilibrium with the aspiration pressure set by fluid level in the waste repertoire is reached (Fig. 2B).
  • the microcarriers form a paste-like material, contained in the transfer tip. This material holds in place, such that the loading column can be removed.
  • the transfer tip is then attached to a syringe and equipped with an injection catheter (Fig. 2C).
  • the microcarrier-based injectable can now be transferred in vivo (Fig. 2D).
  • the drying is composed of 2 separate parts: i) a drying column and ii) a drying box (Fig. 1).
  • the drying column handles the microcarriers, while the drying box is designed to apply a precisely known aspiration pressure and to remove excess pore fluid until equilibrium with the aspiration pressure is reached.
  • the drying columns are themselves composed of 2 parts: the loading column and the transfer tip (as shown in Fig. 2).
  • the loading column is adapted to receive an important amount of volume, typically the content of a 6-well plate well.
  • the transfer tip is plugged at the bottom part of the loading column. It acts as an adaptor for both the column and the needle.
  • the design of the transfer tip must be such that it maintains the compacted microcarriers when lifting the transfer tip off the capillary conductor (or filter membrane).
  • the transfer tip is designed to be conical (cone opening angles between 0° and 180°, preferentially 1° to 90°, more preferentially between 2° and 45°, even more preferentially between 3° and 30°, and most preferentially between 5° and 20°).
  • conical cone opening angles between 0° and 180°, preferentially 1° to 90°, more preferentially between 2° and 45°, even more preferentially between 3° and 30°, and most preferentially between 5° and 20°.
  • Alternative measures to ensure maintenance of the compacted material can be taken, such as temporary placement of a filter membrane to be removed horizontally with minimal force on the microcarrier.
  • the tip opening is also a crucial element for regulation of the fluid flow rate.
  • the tip opening diameter is in between about 20% of the fully expanded microcarrier size to a maximum of about 5cm, more preferentially between 50% of the average fully expanded microcarrier size and 1cm, and even more preferentially between 100% of the average fully expanded microcarrier size and 0.5cm.
  • the box is composed of a drying platform and a waste reservoir.
  • the capillary conductor (contacting cloth) ensures the connection between these two compartments.
  • the capillary conductor is optionally also active as a filter to maintain the particles in the transfer tip during dehydration.
  • its pore size show be adapted to exclude particles from entering the capillary conductor; depending on particle elasticity, it is sufficient to have a pore size smaller than the microcarrier size, or it is necessary to have it much smaller.
  • a filter membrane is imposed between capillary conductor and transfer tip; in this case, the minimal requirement on the capillary conductor is merely not to lose all of its pore water at the desired set pressure to maintain sufficient hydrolic conductivity for the dehydration process.
  • the filter membrane in that case should have a pore size at most on the size scale of the microcarriers, preferably much smaller (including down to nanometric sizes such as standard 0.22 micrometers or 0.45 micrometer filter membranes).
  • the waste reservoir is a Falcon tube into which a whole at a given height has been drilled to ensure a constant level of saline buffer whatever the exact amount of excess medium in the CCM solution.
  • the excess of liquid will indeed leak out of the tube by the hole performed at a specific height defined by the experimental needs. Indeed, polymer concentration may vary from one application to another.
  • a drying box with a preset height of hole at 2cm equivalent to about 200 Pa aspiration pressure), which we find to produce a CCM concentrate suitable for a subcutaneous injection while preserving cell viability.
  • Fig. 3 shows the technical dimensions of a drying column, which consists of two mandatory pieces: the loading column and the transfer tip. Optionally, it can be closed with a cap on its top.
  • Fig. 5 shows the relationship between the dry weight polymer concentration of the biomaterial and the aspiration pressure as set by the height difference Ah between the shelf of the pipet box and the lower edge of the pressure regulator hole.
  • Ah the height difference between the shelf of the pipet box and the lower edge of the pressure regulator hole.
  • the binder fluid in some applications, and in particular to maintain the fluid associated with the microcarriers during extrusion or injection, it is advantageous to render the binder fluid more viscous, for example by adding methylcellulose, carboxymethylcellulose, agarose, starch, polyethylene glycol or any of a large number of cell compatible vicosing agents.
  • methylcellulose, carboxymethylcellulose, agarose, starch, polyethylene glycol or any of a large number of cell compatible vicosing agents Typically, one wishes to increase the viscosity of the interstitial fluid by a factor of 2x to lOOO’OOOx in such applications. For the lower viscosities, addition to cell culture medium may be sufficient.
  • the aim of this example is to estimate the admissible linear compression in sample microcarrier system for co-culture of hematopoietic stem cells and the stromal OP-9 line.
  • Compressible carboxymethylcellulose scaffolds in accordance with W02017/029633 are produced by cryogel bulk scaffold synthesis.
  • a reaction mix consisting of 13.56mg/mL carboxymethylcellulose (AQUALON CMC 7LF PH, 90.5 KDa, DS: 0.84) and 0.486mg/mL adipic acid dihydrazide, buffered with 6.3mg/mL PIPES neutralized to pH 6.7 by 1.2mg/mL NaOH was prepared and filtered through a 0.22um filter (Stericup).
  • the mix was frozen at -20°C in 30mL syringes. After 48h of ciyo-incubation, the syringes are thawed.
  • the scaffolds thus obtained are then fragmented to irregular, compressible microcarriers by extrusion through a 22G catheter. This yields a suspension of microscaffolds suitable for use as microcarriers. These microcarriers are then extensively washed, and autoclaved for sterilization.
  • the surface of the sterile microcarriers is covalently modified with collagen type 1 (from bovine skin, Sigma, C4243), producing collagen- coated carboxymethylcellulose microcarriers (CCM).
  • collagen type 1 from bovine skin, Sigma, C4243
  • CCM carboxymethylcellulose microcarriers
  • the collagen-coated carboxymethylcellulose microscaffolds (CCM) were rinsed twice with D1 water in order to remove the excess of non-adsorbed proteins.
  • covalent crosslinking of the collagen was performed by immersing the CCMs in a solution containing EDC (1 mg/ml) and MES buffer (pH 4.5, 100 mM) in D1 water for 10 minutes.
  • EDC EDC
  • MES buffer pH 4.5, 100 mM
  • Example 3 Microcarrier culture and viability
  • the aim of this example is then to evaluate stability of the collagen-coated carboxymethylcellulose microscaffolds CCM as described in Example 2, seeded with cocultures of murine stromal OP9 and hematopoietic stem and progenitor cells (murine hematopoietic stem and progenitor cells or HSPC, obtained by sorting bone marrow extract for ckit + , lin-, Sca-1 + “KLS”). Seeded at an initial OP9:KLS ratio of 100:1, the CCMs were cultivated for 3 months to achieve steady-state cellular composition prior to mechanical and cellular stability testing against controlled uniaxial compression. This data allows to follow the fate of the co-cultures on the CCMs during controlled compression, and thus to estimate the allowable compression during injection.
  • collagen-coated CCMs were seeded using 75,000 OP9 (GFP) and 7,500 KLS (DsRed) per mg of dry scaffold weight. After initial incubation to allow for cell adherence, the scaffolds were distributed in ultra-low adhesion plates; a 6-well plate (2mg of dry scaffold/well) and a 24 well plate (0.5mg of dry scaffold per well). Culture was performed by adding first 3mL respectively 0.5mL of conditioned medium, and then another 3mL respectively 0.5mL at day 7, followed by half-media changes every week for 3 months. We took care to avoid aspiration of the CCMs by holding the plate in a slanted position for about 30s prior to medium aspiration. This allows the CCMs to sediment and be protected from aspiration. The procedure however is expected to remove a sizeable fraction of the cellular descendance generated by in-vitro hematopoiesis at each medium change.
  • This chamber consisted of a laser cut Perspex sheet in the shape of a microscope slide (i.e. 75mm x 25mm), with a rectangular central observation area (30mm x 12.5mm), closed with a floor consisting of a glued coverslide (Fig. 1A).
  • a laser-cut cleaning cloth made from polypropylene and cellulose (Migros, Switzerland) into the central area (Fig. 6A).
  • the cleaning cloth matched the dimensions of the central observation area (nominally 30mm x 12.5mm, de facto about 0.25mm smaller at each edge due to the finite laser spot size), but in addition we cut a circular hole at its center (8mm nominal diameter, real diameter estimated about 8.5mm).
  • This circular hole serves as sample concentrator, but also as holder for uniaxial compression with a 8mm circular chuck (Fig. 6B). After compression, the samples can be observed by confocal microscope (Fig. 6C).
  • cell culture medium i.e. 50/50 mix of conditioned and fresh medium as described in the main text
  • 50mM HEPES 50mM HEPES
  • Fig. 7 shows the detailed setup used for uniaxial sample compression.
  • the machine is calibrated for force gain by the standard calibration routine in the Texture Exponent user interface.
  • a microscope coverslide identical to the one used for compression chamber fabrication is then placed onto the sample stage of the TextureAnalyzer XTPlus machine, and chuck height calibration is carried out. In this way, the zero-height is defined at upper surface of the coverslide.
  • the compression chamber with previously loaded sample is placed under the chuck, giving rise to the compression setup as shown in Fig. 7.
  • the chuck is then lowered to establish contact with the sample, taking care to avoid air bubbles.
  • the chuck is moved vertically until zero net force is detected; this is done by a built-in routine for this purpose available in the TextureExponent program.
  • the observation chambers were kept closed with a second coverslide during transport and microscopic observation; between experiments, medium was replenished if necessary by placing a drop of medium (50/50 mix of conditioned and fresh medium with 50mM HEPES as described above) such as to keep the clefts between cleaning cloth and Perspex sheet fully filled, indicating minimal capillary pressure. This was necessary since opening and closing by sliding the closing coverslide (Fig. 1C) removed a small amount of medium sticking to the coverslide.
  • Fig. 9 shows confocal images illustrating the effect of various degress of uniaxial compression (0%, 75%, 100%) on the co-cultures of OP9 (green) and KLS+ descendance (red).
  • Fig. 9A and Fig. 9B show that qualitatively, mere contact with the chuck (Fig. 9A) and even transient uniaxial compression by 75% of the original sample height (Fig. 9B) has only minor effects on the co-cultures, whereas 100% compression kills a large fraction of the cells.
  • Fig. 10 shows a quantitative analysis of the percentage of dead cells compared to the estimated total number of cells. While no significant difference can be detected between the material simply loaded into the observation chamber (“Control: Loading” in Fig. 10) and contact with the compression chuck but no actual compression. We detect a significant increase in the dead cell fraction after the 75% uniaxial compression, although the co-culture still remain largely intact (raise of the dead cell fraction from 7% to 13%). Theoretically complete compression on the other hand leads to the immediate death of more than 50% of the cells.
  • composition of the viable cell fraction during compression is a composition of the viable cell fraction during compression
  • Fig. 11 shows the results regarding the composition of the remaining viable cell fraction.
  • This fraction is composed of distinct green (GFP+, OP9) and red (DsRed+, KLS and descendance) fluorescent cells.
  • Control: Loading We find a slight, but statistically significant decrease of the DsRed+ fraction after the chuck zero-force equilibration procedure (from 53% to 45% on average, P-value of 0.04 associated with equilibiration vs. initial loading in linear regression with both the culture and equilibration vs. initial loading as explanatory variables, Bonferroni correction for a total of three tests against the 0% compression condition). Compression by itself then has little influence on the average DsRed+ fraction, although the variability increases especially for the 100% compression, where cells are killed in entire image areas.
  • the strain e in turn is obtained by relating the compressive chuck displacement Ah to the sample height ho :
  • Eq. 1 and eq. 2 correspond to the so-called « engineering » (Moosbrugger, 2002) stress and strain, since the original sample dimensions are used in their calculation(Moosbrugger, 2002).
  • Fig. 12 shows the stress-strain diagrams for the four compression experiments: to 75%, and then again from original sample height to 100% compression, for the two samples prepared (one from the 6-well culture, one from the 24- well culture).
  • the Young modulus is defined as the change of stress per unit of change of strain: da eq. 4 de
  • Fig. 13 outlines the Young moduli as estimated from eq. 4 plotted against the polymer concentration as estimated by eq. 3. Over a wide range, a 3 rd power law is observed for both samples, in agreement with a similar law reported for bulk ciyogel scaffolds. (Beduer et al., 2015a)
  • the polymer concentrations shown on the x-axis of Fig. 13 should be considered indicative: It is quite difficult to determine the exact polymer concentration reached when loading the polymer chamber. For the small volumes used for the compression measurement, the mere act of transferring to a scale causes pore fluid to be lost. We performed two experiments to define the upper and lower bounds of the polymer concentration after loading.
  • the actual polymer concentration is expected to between the upper and lower bounds thus determined. As a reasonable estimate, we indicate the range midpoint and the upper and lower bounds as extremes, and therefore assume the loading concentration to be 8+/- 2mg/mL.
  • Fig. 13 indicates that the implants have Young moduli in the lower kPa range (1.2+/-0. 6kPa). Also, there is well-respected 3 rd power law for the dependency of Young modulus on the polymer concentration throughout a large part of the polymer concentration range throughout the compression experiments.
  • the 3 rd power reflects the mechanics of a porous elastic structure(Beduer et al., 2015a) where increasing compression leads to loss of pore fluid without major change of the mechanical structure.
  • HSPC hematopoietic stem and progenitor cells
  • Flow cytometry and hematopoietic colony forming assays demonstrate the stromal supportive capacity for in vitro hematopoiesis in the absence of exogenous cytokines.
  • Our approach provides a minimalistic, scalable, biomimetic in vitro model of hematopoiesis in a microcarrier format that preserves the HSPC progenitor function, while being injectable in-vivo without disrupting the cell-cell interactions established in vitro.
  • microcarriers of example 2 We used here the microcarriers of example 2.
  • Collagen CCMs surface coating We coated here the microcarriers of example 2.
  • OP9s mesenchymal stromal cell line
  • the total bone marrow cell pellet was resuspended in 3mL volume and loaded into a magnetic separation cell Sorter (AutoMACS, Miltenyi) to remove all lineage positive (Lin+) cells in suspension.
  • the resulting cells were then blocked for 15 minutes on ice (5pg/ml hlgG; I4506-10MG, Sigma Aldrich), and finally stained for one hour on ice with lineage Streptavidin-PO (1/200), as a conjugate to label any remaining Lin+ cells, as well as c-Kit PE-Cy7 (1/200), Sca-1 APC (1/100).
  • the cell suspension was run through a FACS system (Aria Fusion) and the resulting Lin-, c-Kit+, and Sca-1+ (KLS) cells were sorted into Iscove’s Modified Dulbecco’s Medium (IMDM) + Glutamax, 25mM HEPES (31980022, Life Technologies) supplemented with 10% FBS and 1% P/S.
  • IMDM Modified Dulbecco’s Medium
  • 25mM HEPES 25mM HEPES (31980022, Life Technologies) supplemented with 10% FBS and 1% P/S.
  • 2-3 adult male DsRed+ mice were euthanized to collect approximately 200,000 KLS+ cells in suspension for each experiment.
  • FACS cells were kept on ice for approximately 1-2 hours until co-seeding with OP9s on the scaffold.
  • All cells for co-seeding experiments were cultured in 50% fresh basal media (1MDM + Glutamax 25mM HEPES, 10% FBS, 1% P/S) and 50% conditioned 1MDM media (CM).
  • Conditioned media was obtained by culturing confluent GFP+ OP9s with IMDM media for two days (48 hours), filtering the conditioned medium, and freezing the media for no longer than two months at -20°C. After HSPCs and OP9s were collected in suspension and counted, cells were kept on ice for maximum 1 hour.
  • 3D co-seeding Collagen-coated microscaffolds (CCMs, 13.5 mg/ml in PBS) were dried using a cell strainer in a SteriCup filtration system (C3240), using an autoclave cloth to transmit the capillary pressure. Once dried, the globule of CCMs was transferred to a 6- well ultra-low adhesion plate (Corning, CLS347) using the tip of a 2 mL stereological pipette. For each condition, the two cell types (HSPCs, OP9s) were combined, spun down, and re-suspended in a minimal amount of media (approximately 100 pL in total).
  • HSPCs were seeded at the previously established co-seeding ratios, 1:10 “high” (12,000 HSPCs per well) and 1:100 “low” (1,200 HSPCs per well). If limited by HSPC cell number, the 2D condition was performed with only the low seeding density. Cells were fed at D7, complementary to the 3D culture timeline, with 3mL added and no media removed, then cultured at 37°C and 5% CO2 for 12 days in total. Co-seeding experiments were repeated in at least two separate experiments, with technical triplicates within each experiment.
  • Compression testing was performed to assess the limit of compression compatible with cell survival in example 3.
  • Cells were collected for each condition (high/low; 2D/3D) in three fractions: cells in suspension, cells adherent to the CCMs, and any cells adherent to the ultra-low adhesion plates.
  • media was collected in a 50 mL falcon tube, and cells were washed and collected twice with serum-free media.
  • CCMs adherent fraction cells were detached via enzymatic digestion as follows. CCMs were transferred to 24 well plates with a 1000 um pipette tip and 1 mL of collagenase 1 (17100-017, ThermoFisher Scientific) 0.04% was added per well of CCMs for 25 minutes at 37°C and 5% CO2.
  • CFU methylcellulose colony forming unit
  • each CFU plate was read using a StemVision instrument (Stem Cell Technologies), and total colonies were assessed automatically (StemVision proprietary software) and verified manually on the acquired high-resolution whole-plate images according to colony number, size, and cell distribution (Mcniece et al., 1990).
  • Example 3 After 14 days of in vitro culture, seeded-CCMs in suspension were collected from the well plate and poured into the column of the drying device (Example 1) allowing for the CCMs to settle down into the reservoir and reach the desired interstitial aspiration pressure (ca. 200Pa) and therefore polymer concentration (26 +/- 3 mg/ml, Example 3). This condenses the CCMs into a paste-like material with a Young modulus of 1.2+/-0.6 kPa [Example 3 ), which we find sufficient to sustain a 3D architecture in vivo. Sterile syringes were used to aspirate 0.1 ml of coated scaffold without cells, followed by 0.1 ml of air to ensure separation between them. The reservoir with the sedimented cultured scaffold (50 pL) was connected to a lmL syringe and a 20G flexible catheter (BD Biosciences 381703) was plugged in the other end.
  • BD Biosciences 381703 20G flexible catheter
  • NSG mice were chosen for the experiments as OP9 stromal cells are derived from a mixed genetic background and therefore purely syngeneic transplantation was not possible.
  • anesthesia was induced in NSG mice with 4% isoflurane USP- PPC (Animalcare Ltd).
  • An ophthalmic liquid gel (Viscotears, Alcon) was used to protect the eyes and local isoflurane was reduced to 2%.
  • Mice were placed on a heating pad to keep the temperature constant during intervention, and the back of each mouse was shaved at the area of the injections. Betadine (Mundipharma Medical Company) was spread onto the shaved regions to disinfect the skin.
  • a small orifice was created in the disinfected skin using a 18G needle and the 20G catheter (Tro-Vensite i.v. canula, Troge, Hamburg), connected to the loaded syringe, which was gently inserted subcutaneously about 2 cm from the pierced skin.
  • the 20G catheter Troge, Hamburg
  • two separate injections of 50 pL each were performed subcutaneously on either side of the spine. No sutures were required.
  • the mice were placed back in the cage grouped per condition. The entire preparation of the scaffolds and all the injections were performed under the hood to ensure sterility throughout the whole procedure. Each injection, from start to finish, lasted less than 20 minutes per mouse.
  • Animals were treated with antibiotics in drinking water consisting of 30 mg of Enrofloxacin (300 pL of Baytril 10% ad us. vet, 100 mg/mL, Bayer) and 5 mg of Amoxicillin (100 pL of Amoxi-Mepha 200mg/4mL, Mepha Pharma AG) as well as 500 mg of Paracetamol (Dafalgan®) in a total of 250 mL sterile water for the entire duration of the study and replaced every 7 days. Animals were monitored daily by the researchers, and after two weeks, they were monitored daily by animal care services.
  • NSG immunodeficientmice were euthanized 12 weeks post-injections through inhalation of CO2 (6 minutes). The back was shaved gently to better localize the two implants. The samples were harvested in each mouse being careful to keep some subcutaneous tissue around to study the integration of the scaffolds within the normal tissue. Samples were then fixed for 24 hours in 4% paraformaldehyde at 4°C (10 mL PFA in 15 mL Falcon tube), washed three times with PBS, and embedded in paraffin.
  • Tissues were fixed in paraformaldehyde (PFA), submitted for stepwise dehydration and embedded in paraffin blocks for sectioning at 3-4 pm thickness with a rotary microtome (RM, Leica microsystems) .
  • PFA paraformaldehyde
  • RM rotary microtome
  • sections were mounted on glass slides (Superfrost+ slides, Menzel glaser). Paraffin sections were stained with Hematoxylin and eosine (H&E) using the Tissue-Tek Prisma automate (Sakura) and permanently mounted using the Tissue-Tek glas G2-coverslipper (Sakura) to assess morphology.
  • H&E Hematoxylin and eosine
  • Detection of rabbit anti-GFP (Abeam, ab6673, diluted 1:400), rabbit anti- Dsred (MBL, PM005, diluted 1:500), rat anti-CD31 (clone SZ31, Dianova, DIA-310-M, diluted 1:50), rabbit anti-vWF (Abeam, ab9378, diluted 1:100) or rabbit anti-Perilipin (Abeam, ab3526, diluted 1:200) was performed manually. After quenching with 3% H202 in PBS lx for 10 minutes, a heat pretreatment using 0.1M Tri-Na citrate pH6 was applied at 60°C in a water bath overnight. Primary antibodies were incubated overnight at 4°C.
  • Co-seeded CCMs were kept in 3D culture for 12 days in vitro. At days 1, 4, 7, and 11, a small volume of suspended CCMs were removed and transferred to a deep cavity glass microscope slide (produced in lab for imaging, see Example 3, Fig. 6). Within an hour after transferring CCMs to the imaging chamber, they were imaged at varying magnifications (20X for imaging quantification; 63X for cell morphology) using a Zeiss LSM 700 Inverted Confocal Microscope. As the cells were endogenously labeled for GFP (OP9s) and DsRed (HSPCs), serial imaging was conducted at each time point without significant cellular manipulation.
  • GFP OP9s
  • HSPCs DsRed
  • Images used for quantification were composed of a 25-z-stacked, volume- rendered image. To analyze the data, each fluorescent channel was separated and the compiled volume-rendered image was used. Each image/channel was analyzed using Fiji/lmageJ’s threshold tool, with the resulting quantified fluorescent areas converted to cell numbers by using the mean area per cell, as established by manual identification of a subset of the cells. For each 3D experiment, a total of eight independent CCMs were analyzed from two experiments to plot the relative cell proliferation over time.
  • RESULTS The study aimed to provide a microcarrier co-culture system for convenient and minimally invasive injection of a tissue-like living biomaterial, without disrupting cellular viability and multi-cellular interactions during the injection procedure.
  • stromal OP9s and HSPCs on porous CCM microscaffolds (Fig. 14A).
  • the system self-organized such that the OP9 stroma lined the scaffolds coated with collagen 1 to support the HSPC subpopulations (Fig. 14B), and allowed for in vitro studies in diluted microcarrier suspension cultures.
  • the intact co-cultured CCMs, together with their cellular payload were dehydrated by a custom dehydration device (Fig. 14C, Example 1) and delivered in vivo by subcutaneous syringe- injection (Fig. 14D).
  • our starting biomaterials are highly elastic and porous microscaffolds consisting of crosslinked carboxymethylcellulose.
  • a 3D view based on confocal reconstruction after staining with rhodamine 6G is provided in Fig. 14E.
  • the microscaffolds are designed to be reversibly compressible. This allows for facile exchange of the pore fluid by arbitrary sequences of dehydration and rehydration.
  • To obtain the collagen-coated microscaffolds CCM we made use of such cycles to efficiently and covalently functionalize the microscaffolds with collagen type 1 to provide native stromal cell adhesion motives.
  • Collagen-coated, mesenchymal stromal cell-seeded scaffolds promote hematopoietic cell proliferation over time
  • confocal imaging showed effective spreading of OP9 stromal cells on the CCM scaffolds and attachment of the co-seeded HSPCs to the OP9s, with continuous proliferation of hematopoietic cells within the scaffold over 11 days in culture (Fig. 15A).
  • the stromal cells are essential, since at 24 hours HSPCs failed to adhere to the CCMs in the absence of the OP9 cells (Fig. 15B, Fig. 15C).
  • Fig. 15D HSPC nestling within OP9 stromal cells
  • Fig. 15E colony formation within the matrix after 3-4 days of culture
  • HSPCs and their progeny as identified by the DsRed + GFP- population, consisted nearly exclusively of CD45+ cells, as expected for all cells of the hematopoietic lineage derived from bone marrow HSPCs (Fig. 16B) (Weissman and Shizuru, 2008). These cells, which expressed no lineage markers in the cell surface at seeding (Linage- is part of the KLS definition), consisted at day 12 on a mixture of lineage negative and lineage positive cells (Fig. 16B). Acquisition of major lineage markers thus revealed hematopoietic differentiation within the CCM coculture.
  • CFU assays which quantify the number of oligopotent progenitors able to form functional hematopoietic clonal colonies after a 7-10 day culture in semi-solid, cytokine-rich media.
  • Fig. 16F shows an example of a CFU assay at 7 days.
  • the number of relative colonies is significantly higher in the “low” (1:100) seeding condition (2D: 1093.86 ⁇ 567.88, 3D: 764.69 ⁇ 306.84) as compared to the “high” (1:10) condition (2D: 408.34 ⁇ 126.32, 3D: 369.42 ⁇ 99.46), suggestive of either nutrient competition or a negative paracrine regulation by differentiating hematopoietic cells at higher CFU seeding densities. Compared to controls obtained by direct plating of fresh KLS cells without culture prior to the methylcellulose assay, we observed a maintenance, or minor expansion, of the functional HSPC compartment.
  • cytokines commonly used for in vitro HSPC expansion (e.g. thrombopoietin, stem cell factor, or Fms-related tyrosine kinase 3 ligand, (Costa et al., 2018)).
  • the aim here was indeed to provide a minimalistic system enabling further screening without interference from exogeneous cytokines.
  • Compressible porous scaffolds have previously demonstrated promise in providing an injectable solution to traditional cell-based tissue engineering techniques (Beduer etal., 2015b; Bencherif et al., 2012b).
  • Fig. 17A-D shows the workflow for transitioning from a microcarrier-like suspension culture to a transplantable co-culture biomaterial.
  • the CCMs After a predefined time of in vitro culture as a dilute suspension (Fig. 17A), the CCMs are collected and dehydrated to a controlled level by a device specifically designed for that purpose (Fig. 17B, and detailed information in Example 1).
  • the device applies a pre-set hydrostatic pressure to the CCMs by means of a capillary conductor.
  • the hydrostatic pressure sustained by the biomaterial constituted by the CCM is strongly linked to its concentration, this ensures constant material consistency compatible with regards to injection and hematopoietic niche reconstitution in vivo.
  • Example 1 we set a hydrostatic pressure difference of 0.2kPa, equivalent to a fluid level difference of about 2cm, to concentrate the co-culture biomaterial to 26 +/- 3 mg/mL (Example 1). At this concentration, the material remains easily injectable and matches the kPa range for the vascular part of the bone marrow niche (Bello et al., 2018), as detailed in Example 1.
  • the transfer tip is fitted onto a syringe, and assembled with a catheter (Fig. 17C) for subcutaneous injection (Fig. 17D).
  • Fig. 17C a catheter for subcutaneous injection
  • Fig. 17E we assessed whether the procedure of partial dehydration and passage through the catheter during the injection would be harmful to the cells.
  • hematoxylin and eosin standard histological staining revealed intact scaffold particles (Fig. 18A, D and G).
  • the scaffold was host to diverse cell types.
  • Immunohistochemistry (IHC) against GFP revealed strong persistence of confluent OP9 stroma across the scaffolds shown by the large areas stained in brown (Fig. 18E and H). Such staining was absent from the initially cell-free CCM implants (Fig. 18B), providing evidence for the specificity of the anti-GFP staining.
  • anti-DsRed IHC revealed a positive signal only for scaffolds loaded with both GFP+ OP9s and DsRed+ HSPCs (Fig.
  • murine HSPC/OP9-seeded CCM scaffolds can be implanted in NSG mice to produce highly vascularized structures which retain donor stroma and contain locally active hematopoiesis as well as interspersed adipocytes, features reminiscent of adult marrow (Weiss, 2008).
  • CCMs are easily fabricated with standard equipment (freezer, autoclave for sterilization, laminar flow hood for coating under sterile condition), such that production is easily scalable at affordable costand compatible with Good Manufacturing Practice (GMP) production.
  • standard equipment freezer, autoclave for sterilization, laminar flow hood for coating under sterile condition
  • the CCMs feature covalently bound collagen I (Beduer et al., 2018; Serex et al., 2018). Indeed, among a series of extracellular matrix molecules, collagen type I has been shown to provide the highest proliferation levels with KLS cells (Choi and Harley, 2017). Further, contrary to Matrigel used for generation of hematopoietic ossicles (Bello et al., 2018; Hughes et al., 2010; Reinisch et al., 2017), collagen I is a single protein of defined composition that is amenable to clinical use (Lecarpentier et al., 2018; Salvade et al., 2007).
  • OP9 cells have been shown to direct pluripotentstem cells towards the hematopoietic fate, and also have an ability to maintain engraftable hematopoietic stem cells in in vitro 2D co-cultures systems for up to 2 weeks in the absence of additional cytokines (Naveiras, 2009).
  • the combination of OP9 cells and CCMs was found not only to enable the baseline culture of HSPCs over 12 days in culture, but also to provide for easy and effective implantation of the hematopoietic scaffolds in vivo for a follow-up of 12 weeks in NSG mice.
  • the HSPC compartment as measured by the CFU counts in the 2D versus 3D cultures (Fig. 16G), is essentially unaffected by this lower output of total hematopoiesis.
  • the 3D cultures are effectively concentrated in a much smaller volume with higher local cytokine levels (Rodling et al., 2017)
  • our results point towards differential regulation of hematopoiesis in our configuration leading to HSPC enrichment, which may be ascribed to the difference balance of the hypothesized “vascular” versus “endosteal” niches (Leisten etal., 2012; Sanchez-Aguilera and Mendez- Ferrer, 2017).
  • tissue engineered BM has provided a first indicator of supporting subcutaneous, extramedullary hematopoiesis in healthy adult murine tissue without simultaneous ossification. We observe some induction of adipogenesis, pointing towards bidirectional communication between the niche and the hematopoietic compartment.
  • cytodex3 dehydrated material The performance of the cytodex3 dehydrated material is slightly less good (Fig. 20): As for the CCM material, the dehydration device performs perfectly, but a significant amount of cells (OP9) are lost during injection. This is truly cell loss, as the amount of cell death on the cytodex beads does not increase (Fig. 20). Even so, the amount of cells conserved as measured by the GFP area confluence remains significantly above 0, demonstrating that the general method of dehydration and injection can clearly still be used with cytodex 3 microcarriers.
  • Example 6 Cell survival in-vivo in different materials
  • Fig. 21 compares relative in-vivo transfer efficiency for different materials. From Fig. 20, we anticipated some cell loss for microcarriers that are not specifically adapted for the purpose and therefore seeded lOx more cells than on the CCM carriers (Example 2).
  • Fig. 21 The preliminary results shown in Fig. 21 indicate that there is about an order of magnitude of cell loss for the cultisphere G, cytodex 3 and also a non-porous CCM analog. Nevertheless, sufficient cells survive to deliver a living implant, indicating applicability of the method to different microcarriers even though it is evident that engineering appropriate properties into the microcarriers themselves increases the efficiency of the process.
  • OP9 show long-term engraftment in NSG mice (3 months; see also Fig. 18).
  • This example provides characterization of the dehydration box, transfer tip and drying column as described in Example 1, for dehydration of CCM microcarriers as synthesized in Example 2.
  • These data show that for a wide range of set target aspiration pressures (indicated as cm H20 where 1cm H20 is about 98 Pa), the volumetric flow rate is independent on the applied pressure (Fig. 22A).
  • cm H20 where 1cm H20 is about 98 Pa
  • Fig. 22A For low microcarrier concentration as typical for dilute suspension culture, there is also relatively little variation of the flow rate with varying polymer concentration, although the flow rate will decrease at higher polymer concentration when approaching the equilibrium.
  • Example 8 Layered compacted microcarrier culture.
  • FIG. AB shows the layered compacted microcarrier culture.
  • HFF human foreskin fibroblasts
  • Example 9 Integrity of the microcarrier system after exogenous or endogenous reloading
  • This example provides evidence for the integrity of the microcarrier system, once injected, to accept reloading at a second timepoint, whether by reinjection or by colonization of local resident cells that acquire the capacity to colonize the microcarrier through the presence of the stromal layer.
  • a total of 0.4 ml of the partially dehydrated CCM microcarrier suspension was delivered subcutaneously as described extensively in examples 3 and 4.
  • a suspension of 40 microliters of cell-emulating low diffusion carbon microparticles was carefully injected (Fig. 25A).
  • Fig 25B the microcarrier system preserved its geometry after receiving the secondary graft.
  • Fig 25c shows how the cell-emulating microparticles delivered at a second timepoint could distribute through the previously injected microcarrier while preserving the microscopic architecture.
  • Fig. 25D-E 30-50 microliters of the partially dehydrated CCM microcarrier suspension were injected intra-bone, through the murine tibial plate, into the bone marrow cavity.
  • Fig. 25 D-E show at low and high magnification the microcarrier suspension integrating into the bone marrow hematopoietic tissue and thus the potential of the so-delivered microcarrier suspension to incorporate the functional progenitor populations in situ.
  • Luhmes organoids were commercially obtained from Neurix SA, Geneva, Switzerland. They were layered below microcarriers functionalized with Matrigel and then subjected to partial dehydration and culture at a negative pressure of -0.9kPa on a diffusive membrane, while providing fresh medium at a flow rate of 100 micrometers/s on the other side of the membrane. These culture conditions were maintained for 7 days, before fixing the samples with 4% PFA, and processing for paraffine embedding. The samples were stained for Bill tubulin and DAP1. The results of the experiment (Fig. 24 C) show nearly complete integration of the microcarriers into the organoids (80-90% integrated area compared to original size of organoid) with maintenance of organoid viability.
  • Fig. 1 Schematic representation of the hydrostatic pressure applied by the device.
  • a capillary conductor drives the pressure applied by the level in the waste reservoir of the drying box to the sample in the drying column.
  • the pressure can easily be adapted by changing the level of liquid in the waste reservoir.
  • Fig. 2 Overview of the transition from in vitro culture to in vivo injection.
  • A) In vitro co culture of OP9 (stromal cells) and HSPC (KLS) inside our collagen-coated CCM. The biomaterial is used as a microcarrier in a diluted form.
  • B and C) Concentration the diluted CCM by the drying device. The protocol is detailed below.
  • Fig. 3 Technical details of the drying column.
  • the drying column consists of a loading column and a transfer tip.
  • the transfer tip is reversible attached to the loading column by pressing.
  • the loading column can be closed by a cap.
  • Fig. 4 Technical layout of the drying box.
  • the drying box a customized pipet tip box: a waste reservoir with a lateral hole for pressure regulation is inserted into tip shelf, and a capillary conductor in the form of specifically cut wiping cloth is laid out onto the plateau. An extended part of the capillary conductor extends into the reservoir to transmit the aspiration pressure set by the position of the pressure regulator hole.
  • the pipet box also serves as a containment for the excess liquid leaving the waste reservoir by the pressure regulator hole.
  • Fig. 5 Relation between aspiration pressure and the dry weight polymer content of the CCM biomaterial (uncoated precursor). The aspiration pressure is expressed by the height difference between the pressure regulator hole and the plateau of the drying box (Ah in Fig. 4).
  • Fig. 6. Observation chamber adapted for uniaxial compression.
  • sample in the sample area can be compressed uniaxially. Excess liquid is taken up by the cleaning cloth. C) After compression, the sample can be observed by confocal microscopy. For this purpose, the chamber can be closed reversibly by a second microscope coverslide to limit evaporation.
  • Fig. 7 Uniaxial sample compression. A sample is loaded into the circular sample area (Fig. 6, here in cross-section), defined within the cleaning cloth. The compression chamber is then placed onto the sample stage of a mechanical testing machine (TextureAnalyzer XTPlus) and then compressed by uniaxial movement of the chuck. Excess liquid squeezed out from the pore space during compression is drained by the cleaning cloth.
  • a mechanical testing machine TextureAnalyzer XTPlus
  • Fig. 8 Z-projection (maximum intensity) of a confocal stack of a co-culture of OP9 and KLS on CCMs at 3 months. Hoechst 328 was used to stain for nuclei of dead cells.
  • Fig. 9 Z-projection (maximum intensity) of a confocal stack of co-cultures of OP9 and KLS on CCMs after uniaxial compression by 0% (A), 75% (B) and 100% (C) of the original sample height. Hoechst 328 was used to stain for nuclei of dead cells.
  • Fig. 10 Quantification of cell death at different levels of compression.
  • the loading control corresponds to transfer to the compression chamber only, without contact with the chuck; this is closest to the actual culture conditions.
  • 0% compression indicates contact and zero-force equilibration with the chuck of the uniaxial testing machine (TextureAnalyzer XTPlus), but no actual sample compression.
  • 75% indicates a decrease in sample height by 75%, whereas 100% indicates theoretical total compression.
  • the fraction of dead cells is the fraction of predominantly blue cells after Hoechst 328 staining, as compared to the total number of cells (predominantly blue + predominantly red + predominantly green).
  • the compression experiments were repeated twice, once with starting material from a 6-well-plate culture, once with starting material from a 24-well- plate culture.
  • Fig. 11 Quantification of the composition of the remaining viable cell fraction.
  • the DsRed+ fraction is the fraction of red fluorescent cells compared to green + red fluorescent cells; conditions are as for Fig. 10, and evaluation based on the same images.
  • Statistical significance was evaluated pairwise by comparison to the 0% compression condition. For this, linear regression with the experiment and condition (0% compression vs. loading control, 75% compression, 100% compression) as explanatory variables and the DsRed+ fraction as the outcome was used, and the P- value associated with the condition evaluated for significance.
  • Fig. 12. Uniaxial compression Stress-Strain diagram. Four measurements were done: to 75% compression, and then to 100% compression on a total of two samples (one reconstituted from each culture condition).
  • Fig. 13 Young modulus as a function of polymer concentration. Eq. 4 is used for the estimation of the Young modulus from the data shown in Fig. 12; eq. 3 is used to estimate the polymer concentration from the strain shown in Fig. 12.
  • the triangles labelled “0% compression” and “75% compression” outline the estimated polymer concentrations at which the viability quantifications for the 0% compression and 75% compression conditions were done; the grey triangle labeled “Implantation” indicates the estimated concentration of the biomaterial for the in-vivo implantation experiments (26mg/mL, see Example 1).
  • Fig. 14 Transplantable bone marrow niche.
  • stromal cells O9 are combined with hematopoietic stem and progenitor cells (HSPC, selected from bone marrow as lineage-, ckit+, Sca-1+ cells); the resulting cell mix is loaded onto collagen- coated carboxymethylcellulose microparticles (CCMs).
  • CCMs carboxymethylcellulose microparticles
  • CCMs For in-vivo implantation, the cell-loaded CCMs are slowly dehydrated to form a paste-like implantable living biomaterial. Both dehydration speed and final dehydration level are carefully controlled.
  • E Structure of a CCM.
  • F CCM (stained by cell impermeant Hoechst dye) along with green fluorescent stroma and red fluorescent hematopoietic compartment.
  • G Assignment of the different areas as scaffold, HSPCs and lineage-committed progenitors, and stromal cells (OP-9). Confocal images are linearly contrast adjusted.
  • Fig. 15 In-vitro co-culture of OP9 MSCs and HSPCs on CCMs.
  • Qualitative observations from imaging demonstrate large-scale structural outline of the seeded CCMs.
  • C OP9 stroma
  • FIG. 16 3D culture outcome compared to 2D controls via flow cytometry and colony forming assays.
  • the stem and progenitor fraction was finally obtained as cKit+ cells within the Lin-CD45+ population (C).
  • C Lin-CD45+ population
  • D Total CD45+ expansion through flow cytometry, identifying 2D and 3D cell proliferation for both the 1:10 and 1:100 seeding densities.
  • E Total CD45+, cKit+ cell expansion for the same conditions, demonstrating closer similarities between the four conditions.
  • F Total colony count after 7 days in methylcellulose medium, after harvesting total cells from the CCMs (after 12 days of in vitro culture).
  • Fig. 17 Implantation of CCM-based co-cultures.
  • A After seeding, the co-cultures can be cultured in-vitro as classical microcarrier suspension cultures.
  • B To prepare an implant, the material is partially dehydrated to by equilibrating to a predefined hydrostatic pressure level (DR), typically on the order of 0.2kPa (ca. 2cm water column). This is done in a specifically designed transfer tip.
  • DR hydrostatic pressure level
  • the transfer tip is attached to a syringe and an implantation catheter (to avoid accidental intravascular injection).
  • D The co-culture biomaterial is injected subcutaneously.
  • E In vitro assessment of injection viability as quantified through GFP+ OP9 MSC and HSPC cell confluence before, immediately after, and 24 hours after injection (biomaterial seeded 1:100 HSPC - OP9, used at day 1 in vitro).
  • F Macroscopic external view of the implant in the subcutaneous dermal tissue 12-weeks post-implantation.
  • G Visibly vascularized scaffold after sacrifice, seen from the inside of the skin flap.
  • Fig. 18 Histology and cellular composition of implanted scaffolds. Unseeded CCMs (A-C), as well as CCMs cultured with OP9 (D-F) or with 1:10 “high” co- cultures of OP9 and KLS cells (G-I) were implanted into the dorsal skin of NSG mice, and retrieved after sacrifice at 12 weeks. Samples were processed for hematoxylin/eosin (H&E) staining (A, D, G), as well as immunohistochemistry with primary antibodies directed against GFP (B, E, H, marker for OP9 cells) and DsRed (C, F, I, marker of HSPC and progeny).
  • H&E hematoxylin/eosin staining
  • B, E, H marker for OP9 cells
  • DsRed C, F, I, marker of HSPC and progeny
  • Fig. 19 Immunohistochemistry of scaffolds transplanted in vivo. Unseeded scaffolds (A- C) and HSPC/OP9-seeded scaffolds (D-I) were recovered 12 weeks after subcutaneous transplant. Paraffin sections of scaffolds were stained with anti-vWF (A, D, G, arrows indicate megakaryocytes), anti-Perilipin (B, E, H), and anti- CD31 (C, F, I), and antibodies. Scale bars are lOOpm. CD31+ vessels were quantified (J), error bars indicate mean HSD (p ⁇ 0.003).
  • Fig. 21 In-vivo cell transfer efficiency; Various microcarriers were seeded with OP9 cells (GFP+). For the CCM (porous, synthesis described in Example 2), 3*106 cells/mg were seeded, for the others 30*106 cells/mg to compensate for anticipated losses. Relative cell transfer efficiency was quantified from histological cuts as the number of nuclei associated with GFP+ cells.
  • Fig. 22 Dehydration rate measured as volumetric flow rate in CCM as function of target pressure and initial polymer concentration.
  • Fig. 23 Estimated linear fluid velocity at the neck of the transfer tip (maximum expected linear average fluid velocity, 4mm neck diameter). From the data of Fig. 22.
  • Figure 24 A) Non-layered, compacted microcarrier culture of a non-adhesive cell type (BC1, human neural stem cells). B) Layered compacted microcarrier culture of first adhesive cell type (human foreskin fibroblast) and secondary adhesive cell type (BC1). C) Microcarrier-tissue composite (LUHMES organoids). (1) shows the diffusive membrane while (2) shows the microcarrier-tissue composite.
  • BC1 non-adhesive cell type
  • first adhesive cell type human foreskin fibroblast
  • BC1 secondary adhesive cell type
  • M1 Microcarrier-tissue composite
  • LHMES organoids Microcarrier-tissue composite
  • Figure 25 A-C) In situ re-loading of subcutaneously injected microcarrier system.
  • A) Intact skin after re-loading of the microcarrier system with cell-emulating low diffusion carbon microparticles.
  • C) A magnification of the explanted microcarrier system showing the microscale distribution of the re-injected microparticles (black).
  • D-E Intra-bone injection of the microcarrier system.
  • D) Fractured tibial plate at the site of injection (1) and the delivered microcarrier system (2) into the marrow space (3).

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