KR20240053692A - Matrisome-containing synthetic polymer/natural polymer hydrogel-based 3D scaffold and manufacturing method thereof - Google Patents

Abstract

The present invention relates to a 3D scaffold based on a synthetic polymer/natural polymer hydrogel similar to the bone extracellular matrix. Specifically, by filling an existing synthetic polymer scaffold with a natural polymer hydrogel containing matrisomes, the mechanical properties are maintained while the surrounding area is maintained. The attachment, proliferation, and osteogenic differentiation of bone tissue cells are improved, resulting in faster bone regeneration, making it possible to use it as a stem cell culture and bone graft or substitute material.

Description

Matrisome-containing synthetic polymer/natural polymer hydrogel-based 3D scaffold and manufacturing method thereof}

The present invention relates to a 3D scaffold based on matrisome-containing synthetic polymer/natural polymer hydrogel and a method for manufacturing the same.

Bones are composed of hierarchically segmented structures. It consists mainly of hydroxyapatite contained in many bundles of collagen. Due to the combination of rigid inorganic and flexible organic materials, this arrangement provides a unique combination of outstanding mechanical and biological functions. Therefore, bone remodeling is difficult due to the problem of mimicking the damaged area or the original natural state of the bone. tissue loss. If the damaged area exceeds the critical defect area, bone healing does not occur easily. This level of damage requires a scaffold to support new growth and induce tissue regeneration.

The integration of three-dimensional (3D) structures and multifunctional components allows the fabrication of engineered components that mimic natural tissue for bone tissue repair. This precious property is due to its rheological properties and interaction with the surrounding environment. However, most existing 3D scaffolds consist of only one type of biomaterial. These single-component systems may not be able to actively interact with changes occurring in the surrounding environment or healing tissue. For this reason, bone tissue engineering composite scaffolds must be composed of a balanced structure of hard structures and soft, sophisticated biopolymers.

To satisfy this configuration, a porous scaffold was developed using 3D printing technology using PCL, a biodegradable synthetic polymer. PCL is frequently used in 3D printing due to its ability to mimic complex anatomical features due to its significant mechanical properties and excellent formability. Despite its use in the human body, PCL is inherently hydrophobic and has relatively poor biocompatibility.

Additionally, since PCL must be sintered at high temperatures during the printing process, there is a problem in that it cannot be loaded without damaging biomolecules or living cells. To compensate for these shortcomings, various hydrogels that can provide excellent biocompatibility and high loading efficiency of biocomponents can be filled and used.

Despite these advantages, the composition of hydrogel composite scaffolds is still different from that of natural bone extracellular matrix (ECM).

Accordingly, there is a need for synthetic polymer/hydrogel-based 3D scaffolds similar to bone extracellular matrix.

Bio-nanocomposite Polymer Hydrogels Containing Nanoparticles for Drug Delivery: a Review, I. Gholamali and M. Yadollahi, Regenerative Engineering and Translational Medicine, 2021, 7:129-146

Accordingly, as a result of research to solve the above problems, the present invention fills a synthetic polymer scaffold with a natural polymer hydrogel mixed with matrisome to improve mechanical properties, cell adhesion, and cell proliferation rate while maintaining the inherent properties of the hydrogel. The present invention was completed by developing a 3D scaffold based on synthetic polymer/natural polymer hydrogel similar to the improved bone extracellular matrix.

Therefore, the present invention seeks to provide a 3D scaffold based on a synthetic polymer/natural polymer hydrogel similar to bone extracellular matrix, a method of manufacturing the same, and a use thereof.

In order to solve the above problems, the present invention provides a 3D scaffold based on bone extracellular matrix-like synthetic polymer/natural polymer hydrogel containing matrisome.

In addition, the present invention

Preparing a 3D printed synthetic polymer scaffold;

Preparing a matrisome-fibrinogen solution by mixing matrisome and fibrinogen;

Mixing the collagen solution and the matrisome-fibrinogen solution; and

Step of filling the collagen-fibrin hydrogel mixed with matrisome on the 3D printed synthetic polymer scaffold.

A method for manufacturing a 3D porous composite scaffold comprising a is provided.

Additionally, the present invention can provide a bone substitute or bone filler including the 3D porous composite scaffold.

Therefore, the present invention seeks to provide a 3D scaffold based on a synthetic polymer/natural polymer hydrogel similar to bone extracellular matrix, a method of manufacturing the same, and a use thereof.

The 3D scaffold based on a synthetic polymer/natural polymer hydrogel similar to bone extracellular matrix according to the present invention is capable of producing a scaffold with a shape that exactly matches the affected area where bone damage has occurred through a 3D printer, and bonds to the affected area faster than existing scaffolds. And the adhesion, proliferation, and osteogenic differentiation of surrounding bone tissue cells are improved, so a faster bone regeneration effect can be expected. In addition, since steps such as cell transplantation are not required, faster recovery can be expected by transplanting the scaffold to a patient with bone damage in a short period of time.

Figure 1 shows the characterization of the 3D printed scaffold.
Figure 1A shows a schematic diagram of a 3D printed PCL scaffold filled with collagen/fibrin (CF) hydrogel and CF/matrisome (CF/MTS).
Figure 1B shows the mechanical properties of 3D printed PCL scaffolds filled with collagen/fibrin (CF) hydrogel and CF/matrisome (CF/MTS), and
Figure 1c shows PCL, PCL/CF and PCL/CF/MTS by cross-sectional FE-SEM [scale bar; 1mm and 200um].
Figure 2 shows the results of proteomics analysis,
Figure 2A is a pie chart showing the distribution of protein coverage in hMSCs-derived matrices;
Figure 2b shows the distribution of quantified proteins.
Figure 3 shows the results of two-dimensional (2D) model evaluation,
The proliferation rate of GFP-hMSCs (Figure 3a) and the fluorescence image of GFP-hMSCs (Figure 3b) on PCL/CF and PCL/CF/MTS scaffolds (n = 6) are shown [scale bar; 1,000 and 300 μm. * indicates a significant difference of p<0.05 and ** indicates a significant difference of p<0.01].
Figure 4 is the result of a three-dimensional (3D) model,
(B) Proliferation rate of GFP-hMSCs on PCL/CF and PCL/CF/MTS scaffolds (n = 6) (Figure 4A) and (B) fluorescence image of GFP-hMSCs (Figure 4B) are shown [scale bar; 1,000 and 300 μm. * indicates a significant difference of p<0.05 and ** indicates a significant difference of p<0.01].
Figure 5 is an evaluation of osteogenic gene expression levels, (a) collagen type 1 (COL1), (b) osteocalcin (OCN), (c) for PCL/CF and PCL/CF/MTS scaffolds (n = 6). ) Expression levels of osteopontin (OPN), (d) bone sialoprotein (BSP), and (e) bone morphogenetic protein 2 (BMP2) are shown [* is significant difference at p<0.05, ** is p Significant difference of <0.01, *** indicates significant difference of p<0.001].
Figure 6 shows the results of bone regeneration in a calvarial defect of critical size.
(a) μCT images of calvarial defects implanted with PCL, PCL/CF, and PCL/CF/MTS scaffolds. Relative BMD results obtained from quantitative analysis of (b) bone volume, (c) BV/TV, and (d) μCT scanning are shown [* indicates significant difference at p < 0.05, ** indicates significant difference at p < 0.01. indicates].

Hereinafter, the present invention will be described in more detail.

All technical terms used in the present invention, unless otherwise defined, are used with the same meaning as commonly understood by a person skilled in the art in the field related to the present invention. In addition, although preferred methods and samples are described in this specification, similar or equivalent methods are also included in the scope of the present invention. In addition, the numerical values described in this specification are considered to include the meaning of “about” even if not specified. In this specification, when a part “includes” a certain element, this means that it may further include other elements rather than excluding other elements, unless specifically stated to the contrary. The contents of all publications incorporated by reference herein are hereby incorporated by reference in their entirety.

The present invention relates to a 3D porous composite scaffold comprising a biodegradable synthetic polymer and a natural polymer hydrogel containing matrisome.

As used herein, the term "scaffold" is one of the important basic elements in the field of tissue regeneration engineering and refers to a structure or support that serves to provide a suitable environment for cell adhesion, differentiation, and proliferation and differentiation of cells moving from the tissue periphery. do. In particular, it can be used for bone regeneration.

The biodegradable synthetic polymer is polycaprolactone (PCL), polylactide (PLA), polylactide glycolide (PLGA), polyglycolide (PGA), or these. It may be a mixture.

In one embodiment of the present invention, a porous support was fabricated using 3D printing technology using polycaprolactone (PCL), a biodegradable synthetic polymer, which provides a rigid structural framework capable of holding soft components inside.

The term “hydrogel” used herein refers to a three-dimensional structure of hydrophilic natural polymer that retains a sufficient amount of moisture. The hydrophilic natural polymer-based hydrogel can form a composite hydrogel of collagen and fibrin.

In one embodiment of the present invention, a 3D printed PCL scaffold was manufactured by filling it with lagen/fibrin-based hydrogel composed of collagen and fibrinogen. Collagen makes up up to 90% of the organic matrix in bone and has a high water content. High hydrophilicity offsets the disadvantages of PCL and improves cell affinity. Subsequently, fibrin, which is derived from fibrinogen, a natural biopolymer, during the blood coagulation process, is also added to the matrix. This provides improved mechanical strength through rapid coagulation and also aids in tissue repair.

The collagen/fibrin hydrogel is manufactured through a process in which fibrinogen combines with thrombin to chemically undergo primary cross-linking, and collagen undergoes secondary cross-linking.

The physical properties of the hydrogel can be adjusted depending on the volume ratio of collagen and fibrinogen that make up the hydrogel.

The collagen:fibrinogen may have a volume ratio of 1:1 to 1:4. If collagen becomes higher than fibrinogen, there may be a problem in which chemical cross-linking with thrombin does not occur.

It is preferable to dissolve the collagen-fibrinogen complex in phosphate buffered saline solution so that the concentration is 1 to 5%.

In one embodiment of the present invention, Matrisome (MTS) was used mixed with the collagen/fibrin hydrogel.

The term “Matrisome (MTS)” used in the present invention refers to a biomass rich in extracellular matrix (ECM) extracted from human bone-marrow derived mesenchymal stem cells (hMSCs). As a molecule, it can be said to be a complex of proteins and growth factors.

In the present invention, MTS serves to directly supply differentiation factors to bone tissue by loading it into collagen/fibrin hydrogel.

It was confirmed in Figure 1b that there was no change in the physical properties of the support depending on the presence or absence of MTS.

In one embodiment of the present invention, a PCL scaffold can be manufactured by filling the MTS mixed collagen/fibrin hydrogel.

The composite scaffold according to the present invention may include a cross-linking agent in addition to the collagen/fibrin hydrogel, and thrombin or the like may be used as the cross-linking agent.

The composite scaffold may have sufficient strength not to be easily broken during in vivo culture and/or differentiation for 3D cell culture/differentiation or encapsulation. The strength can be adjusted by a person skilled in the art according to known methods depending on the degree of crosslinking of the hydrogel, the concentration of thrombin, the concentration of the hydrogel, etc. Accordingly, if the composite scaffold has sufficient strength, it can be manufactured for cell culture or differentiation purposes capable of 3D cell encapsulation.

In one embodiment, the composite scaffold according to the present invention may have a storage modulus of 500 Pa or more (500-700 Pa).

In one embodiment of the present invention,

Preparing a 3D printed synthetic polymer scaffold;

Preparing a matrisome-fibrinogen solution by mixing matrisome and fibrinogen;

Mixing the collagen solution and the matrisome-fibrinogen solution; and

Step of filling the collagen-fibrin hydrogel mixed with matrisome on the 3D printed synthetic polymer scaffold.

It includes a method of manufacturing a 3D porous composite scaffold comprising a.

The information about the 3D porous composite scaffold described above can be applied to the manufacturing method of the 3D porous composite scaffold of the present invention.

In a specific embodiment of the present invention, the PCL/CF/MTS composite scaffold was observed with a scanning electron microscope (SEM) to confirm that it was manufactured to have pores and interconnections (Figure 1c), and compared to PCL/CF, It was confirmed that the cell proliferation rate was excellent (Figures 3 and 4).

In addition, in another specific embodiment of the present invention, as a result of in vitro analysis of osteogenic differentiation induction on the PCL/CF/MTS composite scaffold, expression of osteogenic differentiation-related proteins such as COL1, OCN, OPN, BSP, and BMP2 This increase was confirmed to confirm that the osteogenic differentiation induction function was excellent (Figure 5).

In addition, in another specific example of the present invention, it was confirmed that the PCL/CF/MTS composite scaffold was implanted in the mouse skull defect area in vivo, and bone formation in the defect lesion area was noticeable (Figure 6 ).

In conclusion, the 3D porous composite scaffold according to the present invention is a composite scaffold manufactured by filling a PCL scaffold with collagen/fibrin hydrogel to which MTS has been applied. It is not only used for stem cell culture but also has a bone regenerative effect and can be used to bone bone during transplantation. Since it can improve cell adhesion and promote bone differentiation, it is expected to be useful as a scaffold for bone regeneration.

In the present invention, the term "bone regeneration" may refer to any phenomenon that treats, alleviates, or improves bone damage through processes such as proliferation of bone cells in damaged bone tissue or differentiation of osteoblasts into bone cells. Although not limited thereto, in particular, the bone regeneration may be achieved by promoting osteogenic differentiation.

Accordingly, the present invention can provide a bone substitute or bone filler including the three-dimensional porous scaffold.

The three-dimensional porous scaffold according to the present invention improves mechanical properties, cell adhesion, and cell proliferation rate while maintaining the inherent properties of hydrogel, and is used as a bone substitute or bone filler, that is, a porous artificial graft material and an artificial graft material for autogenous bone replacement. It can be widely used.

[Example]

Hereinafter, the present invention will be described in more detail through examples according to the present invention and comparative examples not according to the present invention, but the scope of the present invention is not limited by the examples presented below.

[Experimental process]

<Acquisition of materials>

Polycaprolactone (PCL) (MW 45,000), fibrinogen, thrombin, and acetic acid were purchased from Sigma-Aldrich (St. Louis, MO, USA). Collagen (Atellocollagen) was purchased from Dalim Tissen Co., Ltd. (Seoul, South Korea). HumaTein™ (Matrisome, MTS) was purchased from ROKIT Healthcare (Seoul, Korea). DPBS (Dulbecco's phosphate buffered saline) and trypsin-EDTA were purchased from Gibco BRL (Invitrogen Co. Ltd., CA, USA). Cell counting kit-8 (CCK-8) was purchased from Dojindo (Kumamoto, Japan).

MTS was extracted from human bone marrow mesenchymal stem cells and purchased from Lonza (Basel, Switzerland).

<Design and fabrication of 3D printed PCL scaffold>

3D printed PCL scaffolds can be prepared using PCL scaffold methods known in the art (e.g., DN Heo, NJ Castro, S.-J. Lee, H. Noh, W. Zhu and LG Zhang, Enhanced bone tissue regeneration using a It was designed through CAD (Computer-Aided Design) based on 3D printed microstructure incorporated with a hybrid nano hydrogel, Nanoscale , 2017, 9 , 5055-5062. To provide mechanical strength, 3D printed PCL scaffolds were fabricated using an FDM 3D printer (Dr. In vivo, ROCKIT Healthcare, Korea). .

PCL was placed in an extrusion syringe and heated at 85°C to melt the polymer. PCL was dispensed at 700-800 kPa with an extrusion speed of 5 mm s-1.

<Production of collagen/fibrin (CF) hydrogel composite PCL scaffold (PCL/CF)>

First, the collagen-fibrinogen solution and cross-linking solution were prepared separately (presented based on 500 ul).

Collagen stock solution was prepared by mixing 125ul of 9mg/mL acidic collagen solution extracted from rat tail (50%), 25ul of 10 Dulbecco's phosphate buffered saline (DPBS), and 75ul of deionized water. 125ul of fibrinogen solution was prepared at 20mg/mL. Afterwards, the collagen stock solution and the fibrinogen stock solution were mixed to prepare 350 ul of collagen-fibrinogen solution. (Physical cross-linking occurs as collagen's triple helix structure unravels above 15 ℃, so when making collagen stock, it is prepared at 4 ℃)

The cross-linking solution was prepared at 150ul by mixing 25ul of 0.2N sodium hydroxide (NaOH) and 125ul of 6 unit/mL thrombin solution.

150ul of the cross-linking solution mixed well with 350ul of the collagen-fibrinogen solution was sprinkled on the already prepared PCL scaffold, and the sample was quickly transferred to a cell culture dish and incubated at 37°C for 30 minutes for complete polymerization.

<Matrisome-containing collagen/fibrin hydrogel (CF/MTS) production>

Collagen stock solution was prepared by mixing 125ul of 9mg/mL acidic collagen solution extracted from rat tail (50%), 25ul of 10 Dulbecco's phosphate buffered saline (DPBS), and 75ul of deionized water.

After dissolving Humatein powder (MTS) at a concentration of 4mg/ml, 20mg of fibrinogen was added and the MTS solution was mixed to prepare 125ul of MTS-fibrinogen solution. Afterwards, the collagen stock solution and the MTS-fibrinogen stock solution were mixed at a volume ratio of 1:1 to prepare 350ul of the collagen-MTS-fibrinogen solution. (Produced at 4℃)

The cross-linking solution was prepared at 150ul by mixing 25ul of 0.2N sodium hydroxide (NaOH) and 125ul of 6 unit/mL thrombin solution.

150ul of the cross-linking solution was mixed well and sprinkled on 350ul of the collagen-fibrinogen solution, and the sample was quickly transferred to a cell culture dish and incubated at 37°C for 30 minutes for complete polymerization.

<Production of CF/MTS composite PCL scaffold (PCL/CF/MTS)>

Collagen stock solution was prepared by mixing 125ul of 9mg/mL acidic collagen solution extracted from rat tail (50%), 25ul of 10 Dulbecco's phosphate buffered saline (DPBS), and 75ul of deionized water.

After dissolving Humatein powder (MTS) at a concentration of 4mg/ml, 20mg of fibrinogen was added and the solution was mixed to prepare 125ul of MTS-fibrinogen solution. Afterwards, the collagen stock solution and the MTS-fibrinogen stock solution were mixed to prepare 350ul of the collagen-MTS-fibrinogen solution. (Produced at 4℃)

The cross-linking solution was prepared at 150ul by mixing 25ul of 0.2N sodium hydroxide (NaOH) and 125ul of 6 unit/mL thrombin solution.

150ul of the cross-linking solution mixed well with 350ul of the collagen-fibrinogen solution was sprinkled on the already prepared PCL scaffold, and the sample was quickly transferred to a cell culture dish and incubated at 37°C for 30 minutes for complete polymerization.

<Mechanical Characteristics Analysis>

Elastic and viscous moduli SJ Lee, H. Nah, DN Heo, K.-H. Kim, J.M. Seok, M. Heo, H.-J. Measurements were made using a rotational rheometer (MCR 92, Anton-Paar, Austria) according to Moon, D. Lee, JS Lee and SY An, Carbon , 2020, 168 , 264-277.

All samples were prepared in disk shape and measured using a rotational rheometer with a plate-plate geometry with a diameter of 8 mm and a gap of 1 mm to investigate the frequency-dependent viscoelastic behavior. Frequency sweeps were performed in the range 0.1–4.0 Hz at a constant strain of 1%. Storage modulus (G') and loss modulus (G') are: Lee, J. S., Nah, H., Moon, H. J., Lee, S. J., Heo, D. N., & Kwon, I. K. (2020). Controllable delivery system: A temperature and pH-responsive injectable hydrogel from succinylated chitosan. Measurements were made as in Applied Surface Science, 528, 146812.

<Morphological analysis>

The morphologies of PCL, PCL/CF, and PCL/CF/MTS were observed by field emission scanning electron microscopy (FE-SEM; SU8010, Hitachi, Japan) at an accelerating voltage of 5 kV. All freeze-dried samples were taken from a cross-sectional layer cut across the PCL structure and then sputter-coated with platinum (30 mA) for 150 s using an MC1000 ion sputter coater (Hitachi, Japan).

<Antibody array analysis>

The experimental process using Human L-2000 is briefly explained as follows. This array kit can simultaneously detect 2,000 cytokines, chemokines, adipokines, growth factors, angiogenic factors, proteases, soluble receptors, soluble adhesion molecules and other proteins. The experiment was performed according to the manufacturer's instructions. Briefly, antibody array slides (RayBiotech, Norcross, GA, USA) were dried and incubated with blocking solution for 30 min. After decanting the blocking buffer, the slides were incubated with 400 uL of the diluted sample for 2 hours at room temperature. After decanting, each array slide was washed with 800 uL of washing buffer three times while shaking.

Slides were incubated with biotin-conjugated anti-cytokine antibodies for 2 hours at RT with gentle shaking. Then, the slides were washed and incubated with Cy3-Conjugated Streptavidin for 2 hours at room temperature. Finally, the slides were washed, rinsed with deionized water, and centrifuged at 1000 rpm for 3 minutes to remove water. Then, the detection mixture was added and a chemiluminescence imaging system (Bio-Rad Laboratories, Hercules, CA, USA) was applied to detect and visualize the signal. The original signal intensity of the protein was determined by densitometry and normalized using a positive control after correction. Relative protein expression levels were obtained by comparing normalized values.

<Cell culture>

Human bone marrow-derived mesenchymal stem cells (hMSC) (Lonza, PT-2501) and human bone marrow-derived mesenchymal stem cells with GFP (GFP-hMSC) (Cyagen, HUXMA-01101) were used at passage 6 for all cell experiments. Both hMSCs and GFP-hMSCs were grown in mesenchymal stem cell basal medium (Lonza, PT-3238) and a supplement kit (Lonza, PT-4105) containing fetal bovine serum, L-glutamine, and gentamicin as growth medium (GM). Cultured. hMSCs were incubated in a humidified 5% CO ₂ incubator (Thermo, Heracell VIOS 250i). Fresh medium was replaced every 2 days during the cell culture period.

<Cellular behavior of hMSCs in 2D/3D>

hMSCs were cultured in GM and encapsulated within CF hydrogel at a density of cells/mL. After 1, 4, and 7 days, samples were washed twice with DPBS and treated with CCK-8 solution according to the manufacturer's instructions. Briefly, the solution was diluted to the ratio of free GM (no supplements). After incubation in an incubator for 2 hours, absorbance was measured with a microplate reader (BioRad, USA) at a wavelength of 450 nm.

To visualize cell proliferation patterns in 2D and 3D during culture, GFP-hMSCs were cultured in GM for 7 days. GFP-hMSCs were cultured on PCL, PCL/CF, or PCL/CF/MTS, and cells were seeded on the scaffold surface or encapsulated in the scaffold. Growing cells were observed and characterized using the EVOS FL Auto Cell Imaging System (Thermo Fisher Scientific, EVOS-M7000).

<Quantitative real-time polymerase chain reaction (qRT-PCR)>

To evaluate osteogenic gene expression in PCL/CF/MTS scaffolds, hMSCs at a cell/mL density were encapsulated in CF or CF/MTS hydrogels and the hMSC Osteogenic Differentiation Medium Bullet Kit (OM; Lonza, PT-3002) was used. They were cultured for 2 and 4 weeks, respectively. Before RNA extraction, the entire sample was washed with DPBS and treated with TRIzol®Reagent (Invitrogen, 15596018) for total RNA isolation. Afterwards, the isolated total RNA was transcribed into cDNA using AccuPower®RT PreMix & Master Mix (Bioneer, Daejeon, Korea). Quantispeed SYBR mix (PhileKorea, Seoul, Korea) was added to the synthesized cDNA and analyzed with primers (Table 1). All results were normalized using GAPDH.

<In vivo transplantation of PCL/CF/MTS composite scaffold>

11-week-old male Sprague-Dawley rats (Koatech, Korea) were divided into four groups: control group (untreated), PCL group, PCL/CF group, and PCL/CF/MTS group. Care and treatment of animals were performed in accordance with the guidelines established by the Institutional Animal Care and Use Committee (IACUC; DGMIF-20030404-01). A midline incision was made over the calvarium and the full-thickness flap was lifted. A critical size cranial defect (5 mm) was created using a trephine bur under sterile saline irrigation. After forming the defect lesion, each scaffold composition was applied to fill the defect area. The incision was sutured in layers with 5-0 chrome gut and 4-0 silk. All animals received a single intramuscular injection of cefazolin (30 mg/kg) 2 days after surgery. Animals were sacrificed 4 and 8 weeks after surgery.

<Microscopic and histological evaluation>

At 4 or 8 weeks after transplantation, mice from each group were sacrificed, and the surgical site was harvested and fixed in 10% neutral buffered formalin. S. J. Lee, H. Nah, D. N. Heo, K.-H. Kim, J.M. Seok, M. Heo, H.-J. Images were imaged by microcomputed tomography (μCT) using a SkyScan 1172 (Sky-Scan) scanner according to Moon, D. Lee, JS Lee and SY An, Carbon , 2020, 168 , 264-277. The degree of bone density according to μCT was quantified using the average gray value and standard deviation of the region of interest. For μCT analysis, new bone volume (mm ³ ), bone volume ratio (bone volume/tissue volume, %), and bone mineral density (mg/cc) were quantified.

After sacrifice, skull samples were fixed in neutral formalin and decalcified by EDTA treatment. Each sample was embedded in paraffin and sliced into sections. Sample sections were stained with hematoxylin and eosin (H&E) and Masson's trichrome for microscopic histological evaluation.

<Statistical analysis>

All data are expressed as mean ± SD. Statistical analysis of the significance of differences between groups was performed with two-way analysis of Bonferroni's multiple comparison test using GraphPad Prism version 7 (GraphPad Software, CA, USA). Differences according to p-values (*p < 0.05, **p < 0.01, ***p < 0.001) were considered statistically significant.

[result]

1. Characterization of PCL, PCL/CF and PCL/CF/MTS

A schematic diagram of the manufacturing process for the production of MTS-blended collagen/fibrin (CF) hydrogel and 3D printed PCL composite scaffolds is shown in Figure 1A.

First, the 3D printed PCL scaffold was used as a mechanical support for the composite scaffold. The PCL scaffold was fabricated through an extrusion 3D printing system to maintain the frame of the entire scaffold. Additionally, the PCL scaffold provides an external frame and internal pore structure that correlates well with the bone tissue structure. However, PCL has low cell affinity and cannot load cells or biomolecules due to the high-temperature sintering process. Additionally, bones inherently have complex properties that include both hard and soft tissues. To more closely mimic the human body inside the frame, the insufficient biocompatibility of PCL was modified by filling the porous structure of the PCL scaffold with CF hydrogel along with MTS to create a PCL/CF/MTS composite scaffold.

Dynamic rheological properties were measured to confirm the mechanical differences between CF hydrogel, PCL/CF, and PCL/CF/MTS scaffolds (Figure 1b).

The scaffolds were gelled to equilibrium and then studied at room temperature. CF hydrogel corresponds to ductility at G'(160.73Pa) and G"(73.00Pa) at the same Angular Frequency of 1 rad/s. Additionally, PCL/CF scaffold has increased G'(660.77Pa) compared to CF hydrogel. ) and G" (109.75 Pa) values are shown, indicating the strength imparted by the PCL scaffold reinforcing the CF hydrogel. However, there was no difference in physical properties when MTS was incorporated into PCL/CF. This did not affect the rigidity due to the small molecular size of MTS. As can be seen in Figure 1c, the similar cross-sections of the internal microporous structures of both PCL/CF and PCL/CF/MTS in the SEM images support this result. As a result of pore measurement, on average, PCL/CF had a porosity of about 28.8±11.77 um, but PCL/CF/MTS had a porosity of 29.56±2.19 um. PCL itself has no pores, so it cannot load cells or bioactive substances. However, when PCL is reinforced with hydrogel, connected pores are confirmed as shown in the data confirmed by SEM, and cells, media, and nutrients can pass through them.

For the PCL-only group, the image consisted only of a support frame with no filling inside. On the other hand, in the PCL/CF and PCL/CF/MTS groups, the gel was well mixed within the entire PCL structure to form an integrated body. Imaging at higher magnification inside the structure shows that the pores are well connected to each other, allowing water and nutrients to move. Additionally, CF can encapsulate additional biomolecules and cells. In the PCL/CF/MTS group, there was no observable difference in internal structure compared to PCL/CF, indicating that the pore size of the PCL structure was not affected by MTS due to the small molecular size.

Then, proteomics was performed to determine the effect of MTS on cell behavior. MTS proteins consisted of two categories: core proteins and MTS-related proteins (Figure 2a). The core protein is mainly composed of collagen and glycoprotein and accounts for approximately 75% of the total mass. This protein performs a variety of functions, including enhancing cell adhesion and providing signaling through integrin binding. Collagen is very important because it supports ECM components (Figure 2b). Several different types of collagen have been discovered. They contribute to the molecular structure and mechanical properties of many tissues due to their high energy storage capacity and minimal elasticity.

In particular, type IV collagen is the most abundant collagen and plays a role in promoting the formation of bead-shaped filaments. Next, fibrillar collagen accounts for the second largest proportion. This type is abundant in bone because it resists stretching. Additionally, glycoproteins (Fibulin 4, Fibrinogen γ) are highly stable due to their multi-domain structure and play a role in binding adhesion molecules that mediate the assembly of ECM protein components. The remainder of the ECM roughly consists of secreted factors and MTS-related proteins. A number of MTS-related proteins include proteins that function to regulate the ECM (HTRA1, MMP-2, LOXL1/2). They form cross-links between fibrils that control the delicate balance between stability and remodeling and provide mechanical strength to the structure. These secreted factors also play a role in regulating stem cell homeostasis through various cellular processes such as control of cell proliferation and differentiation. Table 2 below presents additional information on core substrates and substrate-related proteins. Therefore, MTS extracted from bone marrow stem cells provides a new environment for cells that act to enhance proliferation and differentiation in a direction related to each secreted factor.

2. Comparison of cell viability and osteogenic gene expression in different dimensions of PCL/CF/MTS scaffolds in in vitro model.

To confirm the biocompatibility of the developed composite PCL/CF/MTS scaffolds, 2D (cells seeded on the surface) and 3D (cell encapsulation) cell culture experiments were performed on the scaffolds. When testing implantable biomaterials used in tissue engineering, it is biologically important to consider both dimensions. This is because the microenvironment in which the scaffold will be placed in the actual body consists of the 3D ECM environment of the surrounding tissue that will be in contact with the surface of the scaffold.

2D cell behavior was analyzed to determine aspects of cell surface contact (Figure 3). hMSCs were seeded on the surface of PCL, PCL/CF, and PCL/CF/MTS. On the first day of measurement, the PCL/CF and PCL/CF/MTS groups showed significantly higher proliferation rates than the PCL group. The difference in proliferation was 135% and 159% for PCL/CF and PCL/CF/MTS, respectively (Figure 3a). This upregulation of proliferation was also observed on days 4 and 7. On day 7, the proliferation rates of each PCL, PCL/CF, and PCL/CF/MTS group were observed to be 141%, 259%, and 323% as follows. Each was compared with the PCL group on day 1.

The low increase in cell proliferation in PCL is due to the surface aspect of the PCL scaffold, which has a smaller surface area than PCL/CF to which cells can attach. In this case, 3D printed PCL itself provides a relatively hydrophobic surface whose porosity is not sufficient for good cell affinity to induce cell attachment. Additionally, the open structure of PCL scaffolds (in the absence of filled hydrogel) provides large spaces inside the scaffold, which limits its ability to retain cell clusters for tissue regeneration.

To overcome the limitations of PCL, rather than chemically ROWLF the PCL surface, we decided to induce improved adhesion and proliferation of hMSCs using hydrogels to mimic the natural ECM structure in which real cells survive. The difference between PCL/CF and PCL/CF/MTS also showed a significant difference in the degree of proliferation. PCL/CF showed a proliferation growth rate of 27% between days 1 and 4, while PCL/CF/MTS showed a proliferation growth rate of 71% during the same period. The growth rate gap between the PCL/CF group and the PCL/CF/MTS group widened to 77% and 93% between the 4th and 7th days, respectively. hMSCs showed excellent adhesion and proliferation in all groups, but proliferation was particularly high in the PCL/CF/MTS group (Figure 3).

These results are closely related to the situation of the scaffold surrounding hMSC. The surface area was spatially increased through the surrounding CF hydrogel, thereby increasing space for cell proliferation. Specifically, MTS-related proteins regulate the ECM to form cross-links between fibrils, thereby providing an appropriate structure with mechanical strength. Additionally, hMSCs morphologically maintained their spindle-like morphology on PCL/CF and PCL/CF/MTS scaffolds. As a result of 2D culture on the scaffold, the PCL/CF/MTS group showed significantly higher cell proliferation. These findings showed that MTS derived from bone marrow MSCs act to increase cell proliferation when loaded.

We also focused on cell encapsulation strategies as part of 3D culture conditions. As mentioned above, because 2D cell culture systems do not sufficiently mimic native ECM, CF hydrogel was used as a 3D artificial ECM matrix that can provide a native environment for hMSCs. In this experiment, a comparative evaluation was performed between the PCL/CF and PCL/CF/MTS groups since PCL alone cannot load cells. As shown in Figure 4A, from day 1 of cell loading, the group containing MTS showed approximately 22% higher proliferation than the PCL/CF group. This pattern continued until day 7, when the experiment ended. On day 7, the growth rate in the PCL/CF group and PCL/CF/MTS group increased by 76% and 140%, respectively, compared to the PCL/CF group on day 1.

The difference in daily growth rate was also the same as the 2D culture result. The PCL/CF group showed a growth rate of 29% from day 1 to day 4, while the PCL/CF/MTS group showed a higher growth rate of 59%. Additionally, the growth rates between days 4 and 7 were 50% and 59% for the PCL/CF and PCL/CF/MTS groups, respectively. This showed a higher growth rate in the PCL/CF/MTS group. Interestingly, this increase in proliferation rate showed a more pronounced difference in fluorescence images at day 7 (Figure 4b). The MTS composite group showed MSC-specific fibroblast-like morphology, which significantly affected the degree of cell adhesion. This results in the development of adhesion-related glycoproteins, including fibulin 4 and fibrinogen γ.

This indicates that the MTS mixed hydrogel according to the present invention provides a favorable environment for hMSCs to attach, and thus the PCL/CF/MTS scaffold of the present invention has sufficient potential to encourage the differentiation of hMSCs into mature osteoblasts. there is. Secreted factors from MTS allow it to regulate cellular processes. The environmental conditions under which cells can survive and differentiate are very important and depend not only on dynamic mechanical loads but also on the mechanical properties of the ECM. Additionally, initial cell adhesion is important for tissue regeneration. Cell adhesion can be influenced by both surface and biochemical properties of the scaffold. Due to the similar strength and porosity of PCL/CF and PCL/CF/MTS (Figures 1b–1c), these materials could provide excellent cell compatibility for hMSCs to differentiate. Additionally, these types of in vitro results correlated with the results of the respective proteomic analyzes (Figure 2 and Table 2). The main components of MTS, a collagen and glycoprotein, play an important role in the adhesion of mammalian cells and binding to the ECM. This had a direct impact on the 2D and 3D culture results in our experiments.

Before implanting the PCL/CF/MTS scaffold into the mouse calvarial defect model, osteogenic in vitro analysis was performed through qRT-PCR (FIG. 5). hMSCs were well encapsulated in PCL/CF/MTS at 2 and 4 weeks after culture. Next, we investigated whether PCL/CF/MTS could induce osteogenic differentiation of hMSCs. qRT-PCR showed higher expression levels of osteogenic markers such as COL1, OCN, OPN, BSP, and BMP2. This time-dependent upregulation of osteogenic gene markers was significantly higher at 4 weeks, especially for BSP expression (Figure 5c). ECM expressed by osteoblasts is composed mainly of type 1 collagen, but BSP is the highest among non-collagenous proteins. The high expression level of BSP in the PCL/CF/MTS group can be interpreted as very important in that it indicates bone formation. The higher BSP expression level in the PCL/CF/MTS group can be considered very significant in that it indicates the osteogenic differentiation of hMSCs in the CF hydrogel due to MTS loading compared to the PCL/CF group. These results strongly supported that key components of the MTS allow MSCs to adhere to the CF matrix and provide a sufficient environment for proliferation and osteogenic differentiation.

3. In vivo characterization of new bone tissue formation after PCL/CF/MTS scaffold implantation in a mouse calvarial defect model

To evaluate the bone regenerative effect of PCL/CF/MTS, in vivo implantation of the scaffold was performed in a mouse calvarial defect model. Mice were sacrificed at 4 and 8 weeks after transplantation. As seen in the visualized μCT images, a small amount of bone formation was observed in the PCL group, whereas bone formation was prominent around the defect lesion in other groups, including CF and CF/MTS (Figure 6a). The amount of new bone was quantitatively measured by bone volume (BV, mm3), bone volume per total tissue volume (BV/TV, %), and bone density (BMD, mg/cc) (Figures 6b, 6c, and 6d).

All three groups showed no statistically significant differences during the 4-week results for BV, BV/TV, and BMD assessments. However, the PCL/CF/MTS group showed the numerically highest new bone formation. The regenerative effect of the PCL/CF/MTS group was noticeable in the 8-week results. The BV of the PCL/CF/MTS group was calculated to be 14.82 ± 0.99 mm ^{3 ,} which was significantly higher than that of the PCL group (10.78 ± 1.6 mm ³ ). The results of BV/TV derived from μCT showed more bone growth in the PCL/CF/MTS group (10.38 ± 0.69%) compared to the PCL group (7.55 ± 1.12%) and PCL/CF group (8.86 ± 0.89%). shows similar aspects to the BV value in that (Figure 6c). Additionally, the transplant site BMD in the PCL, PCL/CF, and PCL/CF/MTS groups was 56.8 ± 8.96 mg/cc, 76.53 ± 2.67 mg/cc, and 85.49 ± 9.85 mg/cc, respectively, compared with the PCL group. The regeneration of the MTS group showed significant improvement (Figure 6D).

H&E staining was performed to demonstrate new bone formation in host tissue (Figure 7a). In the PCL group, the transplanted defect area was filled with a small amount of bone and fibrous tissue. However, the PCL/CF and PCL/CF/MTS groups showed new bone with better formation and intimate contact with the host tissue surrounded by fibrous connective tissue. In the PCL/CF/MTS group, new bone formation was observed with several bony islands formed between the fibrous tissues. Additionally, Masson's Trichrome staining was performed to evaluate and visualize new bone formation in the mouse skull defect model (Figure 7b). The PCL scaffold alone was observed to have low biodegradability within the defect area and poor compatibility with host tissue at the implant site. However, the CF and CF/MTS combination groups showed better new bone formation and fibrous tissue surrounding the PCL scaffold. New bone formation in specific areas appears larger in the PCL/CF group, but the dense fibrous tissue around the new bone formation area appears larger in the PCL/CF/MTS group. This result also appeared to be related to the proteomics data (Figure 2).

MTS contains collagen-based molecules, and collagen is a major component of the primary ECM. These results indicate that the PCL/CF/MTS group formed more new fibrous tissue than the other groups. Additionally, a new bone area was observed in the middle of the fibrous tissue in the PCL/CF/MTS group. Forms the skeleton and induces the formation of more new bone. These results show that the MTS composite bioscaffold provided an ECM environment that could induce the formation of a higher amount of bone tissue due to MTS. These μCT and histology results are correlated with in vitro experimental results indicating that MTS incorporating bioscaffolds derived from bone marrow MSCs can accelerate bone regeneration.

[conclusion]

The 3D printed PCL composite scaffold filled with MTS and CF hydrogel according to the present invention demonstrated the ability to enhance both cell behavior and bone healing at the defect site by delivering bone tissue-targeting biomacromolecules required for regeneration. 3D printed PCL scaffolds were used to provide mechanical support to resist physical stresses occurring in bone tissue. MTS was used to impart osteogenic function. This component did not interfere with the original mechanical properties of the scaffold after being incorporated into the CF hydrogel. Additionally, the MTS composite scaffold enhanced hMSCs proliferation and induced upregulation of osteogenic gene markers in both 2D and 3D culture environments. As a result, the PCL/CF/MTS composite scaffold improved bone induction and tissue repair in a critical-size defect mouse calvarial defect model. The PCL/CF/MTS composite scaffold according to the present invention can be used as a bone graft or substitute material for improved bone growth, and is therefore expected to be useful in the field of tissue engineering.

Claims (16)

3D porous composite scaffold comprising biodegradable synthetic polymer and natural polymer hydrogel containing matrisomes.
According to claim 1,
The biodegradable synthetic polymer is polycaprolactone (PCL), polylactide (PLA), polylactide glycolide (PLGA), polyglycolide (PGA), or these. 3D porous composite scaffold as a mixture.
According to claim 1,
The natural polymer hydrogel is a hydrogel mixed with collagen and fibrin, a 3D porous composite scaffold.
According to claim 1,
A 3D porous composite scaffold containing collagen and fibrin in a volume ratio of 1:1 to 4.
According to claim 1,
A 3D porous composite scaffold containing matrisome (MTS) and fibrinogen at a weight ratio of 1:5 to 10.
According to claim 1,
The biodegradable synthetic polymer is a 3D porous composite scaffold with a weight average molecular weight ranging from 10,000 to 100,000.
According to claim 1,
A 3D porous composite scaffold combining collagen/fibrin hydrogel composed of collagen and fibrinogen on a PCL scaffold.
.
Preparing a 3D printed synthetic polymer scaffold;
Preparing a matrisome-fibrinogen solution by mixing matrisome and fibrinogen;
Mixing the collagen solution and the matrisome-fibrinogen solution; and
Step of filling the collagen-fibrin hydrogel mixed with matrisome on the 3D printed synthetic polymer scaffold.
Method for manufacturing a 3D porous composite scaffold comprising.
According to claim 8,
The biodegradable synthetic polymer is polycaprolactone (PCL), polylactide (PLA), polylactide glycolide (PLGA), polyglycolide (PGA), or these. Method for manufacturing a 3D porous composite scaffold that is a mixture.
According to claim 8,
A method of producing a 3D porous composite scaffold comprising collagen and fibrinogen in a volume ratio of 1:1 to 4.
According to claim 8,
A method for producing a 3D porous composite scaffold comprising matrisome and fibrinogen in a volume ratio of 1:5 to 10.
A 3D porous composite scaffold prepared from the manufacturing method according to claim 8.
According to claim 12,
3D porous composite scaffold with storage modulus greater than 500 Pa.
A bone substitute comprising a 3D porous composite scaffold according to claim 1 or claim 12.
A bone filler comprising the 3D porous composite scaffold according to claim 1 or claim 12.
A composition for cell culture comprising the 3D porous composite scaffold according to claim 1 or claim 12.