CN110760076B

CN110760076B - Injectable high-strength composite hydrogel based on colloidal particle-iPRF dual-network structure and preparation method and application thereof

Info

Publication number: CN110760076B
Application number: CN201911076477.7A
Authority: CN
Inventors: 王华楠; 陈楷文
Original assignee: Dalian University of Technology
Current assignee: Dalian University of Technology
Priority date: 2019-11-06
Filing date: 2019-11-06
Publication date: 2022-04-08
Anticipated expiration: 2039-11-06
Also published as: CN110760076A

Abstract

The invention relates to the technical field of biomedical materials, in particular to an injectable high-strength composite hydrogel based on a colloidal particle-iPRF double-network structure, and a preparation method and application thereof. The nano colloidal particles in the double-network structure composite hydrogel are assembled under physical action such as electrostatic interaction to form a first heavy colloidal gel network, the platelet-enriched fibrin iPRF can be injected to form a second heavy fibrin hydrogel network, and the two heavy gel networks are physically crosslinked by electrostatic action, hydrogen bond action and hydrophobic action to form the double-network structure hydrogel. Adding a non-anticoagulant drug into a fresh blood sample, performing low-speed centrifugation operation, taking all yellow liquid on the top layer, quickly blending with colloidal particle dry powder, and completely curing at room temperature or body temperature for no more than 2000 seconds. The double-network structure composite hydrogel is a regenerated ideal injectable and moldable biomedical material.

Description

Injectable high-strength composite hydrogel based on colloidal particle-iPRF dual-network structure and preparation method and application thereof

Technical Field

The invention relates to the technical field of biomedical materials, in particular to an injectable high-strength composite hydrogel based on a colloidal particle-iPRF double-network structure, and a preparation method and application thereof.

Background

Achieving repair of damage to human tissues/organs is a key problem in clinical medicine. However, at present, the clinical treatment of human tissue/organ repair means is mainly based on autologous tissue, allogeneic tissue or xenogeneic tissue. Taking bone repair as an example, the "gold" treatment standard for clinical bone repair is patient autologous bone graft filling. Since autologous tissues/organs are not immune-rejected and usually have good vascularization, therefore, the repair is performedThe compound effect is obvious. For example, the best clinical bone repair protocol for a bone defect is to take a bone graft from an iliac site of a patient's own body to a site of bone defect affection; although the clinical repair effect is good, the defects are obvious: 1) causing secondary trauma and increasing pain for the patient, 2) limited donor and inability to meet the repair requirements for larger bone defects. Allogeneic and even xenogeneic tissue/organ transplantation is therefore a suboptimal option for clinical tissue repair treatment. For example from Stryker, USA

The bone repair filling material is from calf bones. However, such bone repair materials also have a number of problems that are difficult to solve: 1) limited donors, 2) significant autoimmune rejection of the recipient after allogeneic tissue/organ transplantation, drug suppression on the immune system of a patient, various complications and other organ injuries of the patient, 3) potential disease transmission in the allogeneic and xenogeneic tissue transplantation process, and 4) difficult problems of immunogen removal treatment, storage and quality control system construction of xenogeneic and xenogeneic tissues/organs.

In recent years, research and rapid development in the field of regenerative medicine aiming at assisting, repairing and replacing tissue/organ defects caused by human body trauma or diseases are carried out by taking artificial materials, bioengineering technology, cell therapy technology and drug controlled release technology as means, and a new path is opened for repairing and regenerating human tissue organs. The tissue engineering technology is an important branch of the field of regenerative medicine, combines artificial biomaterials, bioactive factors (growth factors, cytokines, polypeptides, siRNA and the like), stem cells with multi-differentiation functions or functional cells differentiated from specific tissues, combines and constructs tissues/organs with functionality by an engineering method, and finally transplants the tissues/organs into a body for tissue repair and regeneration. However, over the course of thirty years, substantial clinical application of tissue engineering techniques remains the phoenix unicorn, with key challenges mainly including: 1) lack artificial biomaterial system of the bionic extracellular microenvironment which replaces the extracellular matrix of human tissue comprehensively, 2) the quantitative production difficulty of bioactive factors, lack of carrier materials capable of slowly controlling the release of protein drug factors, the difficult overcoming of side effects (ectopic osteogenesis, tumor and the like) except the therapeutic effect of active protein, 3) the difficult control of the differentiation in the stem cell body, limited cell sources, the immunological rejection of allogeneic cells, the difficult establishment of a system for extracting and expanding cells to a transplantation treatment period and the clinical quality control.

Biomaterials are the basis of tissue engineering, and their functions are as Extracellular matrix (ECM) and as carriers of biological signal molecules, controlling and regulating the functions and behaviors of cells by continuously releasing these factors into the physiological tissue environment, so as to finally achieve tissue repair. However, the conventional tissue engineering scaffold materials are far from meeting the requirements of tissue repair and reconstruction in the regenerative medicine application process. The key challenges include: 1) the lack of controllable degradation rate of the stent material; 2) the controllable release of biological factor drugs can not be realized, and the research results show that: the growth factors are directly mixed into the block scaffold, so that the burst release is generated at the initial stage of implantation to generate weak effect on tissue repair, or the excessive release generates the proliferation of bone tissues; 3) difficulty in clinical operation, lack of injectability and plasticity, difficulty in Minimally invasive surgery (Minimally invasive surgery); 4) it is difficult to provide an ideal extracellular microenvironment for the loaded cells, and to ensure the activity and functionality of the cells.

The hydrogel is used as a matrix material which is most widely used for researching the cell microenvironment at present, the typical hydrogel is composed of hydrophilic macromolecules, is rich in water similar to natural ECM, allows biological macromolecules to diffuse and spread through a porous gel network, can regulate and control the hardness through the crosslinking degree, is easy to chemically modify, and is one of the most important biological materials in the fields of tissue engineering and drug controlled release. Typical hydrogels can form stable macrostructures and mechanical strength through intermolecular bonding, and the formed porous network wraps cells in three dimensions and provides a platform for the cells to attach, spread, migrate and proliferate. And for the chemical crosslinking hydrogel, the cells can be blended with a hydrogel prepolymer (or monomer) aqueous solution and then the chemical crosslinking/polymerization is initiated to realize the three-dimensional immobilization of the cells. However, the problems still remain: 1) usually, chemically crosslinked hydrogels need to be introduced with chemical reactionsGel curing is performed, and chemical reactions are usually cytotoxic, so that the survival rate of cells after embedding is reduced; 2) for clinical application, although the prepolymer/cell blending aqueous solution has fluidity and can be injected, the chemical reaction in the hydrogel curing process often has obvious cytotoxicity; 3) whereas for physically crosslinked hydrogels (e.g., collagen), because the gel network is formed based on weak physical bonds, the mechanical properties of the resulting hydrogel are poor (e.g., the maximum elastic modulus E of the collagen hydrogel), despite good biocompatibility_max1kPa), difficult to apply clinically; 4) poor drug release capacity, Meidun force

The collagen scaffold material is a common growth factor (BMP-2) loaded collagen scaffold material for bone repair in clinics, but because the collagen scaffold lacks the controlled release capability to protein factors and factor burst release is severe after initial implantation, the loading amount of protein drugs needs to be increased to realize bone induction, so the cost is high and the bone induction performance is limited. Therefore, the development of a novel hydrogel system which has good biocompatibility, better mechanical strength, injectability and plasticity and can realize the sustained and controlled release of the protein factor medicament has revolutionary significance for the clinical tissue organ repair and regeneration.

In the tissue repair process, the tissue repair and regeneration process is subject to several complex biological processes of blood clot formation, inflammation and tissue regeneration, wherein the role of bioactive signal factors which can induce cell aggregation, proliferation and differentiation in the tissue repair process is important. Especially for bone and cartilage tissues with slow growth cycles, the long-term existence of growth factors around the bone defect part can accelerate and catalyze the repair and regeneration of the bone defect. However, the blood clot formed in the early stage of bone defect, although rich in various kinds of angioblasts and osteogenic growth factors, is difficult to continuously provide bioactive factor drugs at the defect site to induce bone growth because the blood clot is degraded and absorbed rapidly.

The expression of the active protein factors in time and space has important induction effect on tissue regeneration. Therefore, the controlled delivery of growth factors is important for tissue engineering to realize tissue repair and regenerationTechnical challenges. For example, a number of growth factors play a regulatory role in bone regeneration, and some growth factors (e.g., BMP-2, BMP-7, VEGF, FGF-2, etc.) have shown great potential for use in preclinical studies. In recent years, much work has been done to bind growth factors to biomaterials, and long-term induction of bone tissue growth through the release of growth factors is expected. Typical examples include bone tissue repair biomaterial carriers based on collagen sponge material as carrier, BMP-2 (bone morphogenetic protein-2, a bone inducing growth factor) and BMP-7, developed by Medtronic and Stryker, respectively, USA. Especially of Medtronic

The product, although having great clinical success in the field of bone repair regeneration, directly dissolves growth factors in water rapidly, and then mixes with preformed collagen scaffold materials, the growth factors are loaded on the surface of the carrier materials only by surface adsorption, the loading efficiency of the growth factors is low, and protein factors are rapidly burst released along with the implanted materials, and the growth factors cannot be released in the repair area in the upper period, so the product needs to stimulate bone induction by the growth factors far higher than the physiological level, the product cost is increased, and the growth factors far higher than the physiological level cause side effects such as ectopic osteogenesis and the like. In addition, when the bioactive protein drug factors are blended with materials, the conformational change inactivation easily occurs, the materials are easy to be subjected to enzymolysis in vivo, a large dose is often needed to achieve the treatment effect, and the toxic and side effects are often caused.

In summary, growth factors have proven to have a number of problems during clinical treatment, including their susceptibility to inactivation in vivo, side effects and safety factors, and thus, successful clinical transformation applications have been difficult to achieve. For example, VEGF has a potent angiogenic effect, but it may also cause systemic hypotension and edema. BMP-2 is degraded and inactivated very quickly in blood, and has reported risks of ectopic osteogenesis and tumorigenesis. More importantly, the growth factor as a bioactive drug has high production cost and has a plurality of application problems in the storage and use processes. Currently, the bottleneck for this therapeutic strategy of release of growth factors by sustained-release systems is: 1. poor sustained and controlled release capacity of the carrier and side effects caused by local application of the growth factor with burst release super-physiological dose; 2. exogenous growth factors are expensive and have unstable therapeutic effects. Therefore, carrier materials for maintaining protein activity and realizing sustained and controlled release of protein drugs are urgently needed.

In view of the above, the use of autologous concentrated blood products represented by platelet-rich plasma (PRP), concentrated growth factor (PRF) or platelet-rich fibrin (CGF) and derivatives thereof as transplantation and replacement materials for bone and other tissue repair has been suggested. The blood product is separated from the blood of a patient by high-speed centrifugation, and the preparation is simple and the cost is low. Meanwhile, the composition contains various high-concentration growth factors and chemotactic factors, is beneficial to tissue repair, and does not generate immunological rejection reaction. However, the therapeutic effect of blood products and their derivatives alone as filling and replacement materials for bone defects is still controversial. Firstly, the blood products have poor mechanical properties and are difficult to achieve the effect of maintaining space; secondly, the gel blood products are degraded and absorbed too fast in vivo, and are difficult to be stably maintained at the defect parts; in addition, the phenomenon of burst release of growth factors is brought about by the excessively rapid degradation and absorption of blood product gels, resulting in side effects which are difficult to predict. All of the above factors may cause uncertainty in prognosis after filling and repairing bone defects. In addition, some scholars add other biomaterials to blood products and their derivatives, such as PRP, in order to enhance their mechanical properties and regulate the release of factors to promote tissue healing. The above strategy indeed enhances mechanical properties and accelerates healing of bone and cartilage. However, the above materials still have the following drawbacks: 1. PRP contains exogenous additives such as anticoagulant and the like, and the regeneration potential of the PRP is limited; 2. the obtained material has poor mechanical strength, no injectability, complex operation and difficult accurate quantification.

In recent years, Choukroun J introduced the concept of low speed centrifugation into PRF formulations, resulting in a novel blood product that is injectable, concentrated in growth factor (iPRF). The novel blood product does not contain any exogenous additive like PRF, and contains a large amount of platelets, fibrin, various growth factors, chemotactic factors and the like. iPRF has been used for regeneration and repair of bone and cartilage tissue due to its injectability and enriched growth factors. However, the material still has the defects of poor mechanical property, fast in-vivo degradation and the like. It is noteworthy that, in the case of low-speed centrifugation, only a small fraction of the platelets rich in iPRF are activated and thus can remain in the sol state for about 15 minutes, with platelet degranulation and thrombin activation, fibrinogen therein undergoes self-crosslinking to form a fibrin-platelet complex in the gel state. This sol-gel phase transition property makes possible the use of iPRF in combination with other biomaterials. However, the iPRF has short degradation period after implantation, and the growth factor component contained in the iPRF for inducing regeneration is also degraded and inactivated quickly, so that the iPRF has good biological activity and regeneration inducing function, but is difficult to realize long-term and controllable factor release and limited regeneration inducing effect. In summary, if the bioactive factor in iPRF is combined with the controlled-release carrier or the inorganic particle material with osteoinductivity, the combination will certainly promote the tissue repair and regeneration.

Disclosure of Invention

In order to overcome the defects in the prior art, the invention provides a high-strength injectable double-network structure composite hydrogel medical material, and a preparation method and application thereof.

In order to realize the purpose, the invention adopts the following technical scheme:

the invention provides a colloidal particle-iPRF double-network structure composite hydrogel, which is formed by injecting platelet-enriched fibrin iPRF based on colloidal particles and a blood extract to form a hydrogel with a double-network microstructure; the hydrogel is characterized in that colloidal particles form a first heavy colloidal hydrogel network with a self-repairing effect under the physical action, injectable platelet-enriched fibrin iPRF forms a second heavy fibrin hydrogel network, and the two heavy hydrogel networks are physically crosslinked by electrostatic action, hydrogen bond action and hydrophobic action to form the hydrogel with a double-network structure.

In the above technical solution, further, the injectable platelet-rich fibrin iPRF is a blood extract from an autologous, allogeneic or xenogeneic source.

In the above technical solution, further, the colloidal particles are assembled by electrostatic interaction, and are specifically formed by combining any one of the following methods:

1) the glass is formed by respectively blending and assembling positively charged silicon dioxide, bioglass, polylactic acid nanoparticles, negatively charged silicon dioxide, bioglass and polylactic acid nanoparticles, wherein the mixing ratio of the positively charged particles to the negatively charged particles is 0.1-10;

2) the composition is composed of one or more of montmorillonite, chitosan, alginic acid, hyaluronic acid, fibroin and gelatin nanoparticles which are locally provided with positive charge and negative charge groups;

in the technical scheme, the size of the colloidal particles is from nanometer to submicron scale, the particle size is 10 nm-2 μm, the surface charge of the colloidal particles is-40 mV, and the colloidal particles can be assembled to form a colloidal gel network with shear thinning and self-repairing performance through electrostatic interaction among the particles.

Preferably, the colloidal particle has a particle size of 10nm to 500 nm.

The invention provides a preparation method of colloidal particle-iPRF double-network structure composite hydrogel, which comprises the following steps:

centrifuging a fresh blood sample for 1-10 min at a rotating speed of 50-1000 g, wherein the centrifuged and layered top-layer yellow transparent liquid is the blood extract injectable platelet-enriched fibrin iPRF, rapidly and fully mixing the upper-layer yellow transparent liquid with colloidal particles, and curing to obtain the colloidal particle-iPRF double-network structure composite hydrogel; or

Adding colloidal particles into a fresh blood sample, placing the sample into a centrifuge for centrifugation for 1-10 min at a rotation speed of 50-1000 g, and obtaining a top-layer brown yellow colloid after centrifugal layering, namely the colloidal particle-iPRF double-network structure composite hydrogel;

wherein the volume fraction phi of the colloid particles is 0.05-1;

the curing time of the colloidal particle-iPRF double-network structure composite hydrogel is 10-2000 seconds, wherein the injectable and shapeable time window is less than 2000 seconds.

In the above technical solution, further, the colloidal particles may be one or a combination of two of a lyophilized powder of the colloidal particles or a dispersion of the colloidal particles in an aqueous solution.

The invention provides an application of the double-network structure composite hydrogel as a carrier of medicines, bioactive proteins and living cells in a tissue engineering scaffold material, wherein the application is used for tissue repair filling.

In particular to a filler for repairing bone tissues, cartilage tissues, muscles and blood vessels.

The fourth aspect of the invention provides an application of the double-network structure composite hydrogel in a three-dimensional cell culture scaffold or a 3D biological printing ink material added in cells, the double-network structure composite hydrogel is used for 3D biological printing of tissues or organs, 1-100 mmol of anticoagulant is added before printing, and the double-network structure composite hydrogel is soaked in Ca-containing Ca after printing²⁺In aqueous solution, a scaffold with a stable structure is obtained.

According to the colloidal particle-iPRF double-network structure composite hydrogel disclosed by the invention, the particle size of gelatin particles is less than 2 microns, particularly the colloidal particles with the diameters of 10nm-500nm are preferred, and strong non-covalent bond actions (such as electrostatic action, hydrogen bond action, hydrophobic action and the like) exist among the particles, so that the gelatin particles can be self-assembled in an aqueous phase solution to form a gel network structure assembled by the colloidal particles and serve as a first heavy network of a double-network hydrogel system; the second network in the double-network hydrogel is a gel network formed by covalent bond crosslinking formed by fibrin in Injectable platelet-rich fibrin (iPRF) of blood extract, and the two networks are also formed by crosslinking through physical actions such as electrostatic action, hydrogen bond action, hydrophobic action and the like.

The invention has the beneficial effects that:

1. the double-network gel structure endows the hydrogel with the effects of reinforcement and toughening:

in the colloidal particle-iPRF double-network structure hydrogel system, colloidal particles can form a continuous colloidal network instead of being only dispersed in a fibrin network of a continuous phase, so that the mechanical strength of the double-network hydrogel is improved by orders of magnitude. The storage modulus of the colloidal particle-IPRF hydrogel is more than 5kPa and is at least 50 times that of the hydrogel formed by iPRF compared with the storage modulus of the hydrogel formed by pure iPRF of about 0.1 kPa; compared with the colloidal hydrogel formed by pure colloidal particles, the storage modulus of the colloidal particle-iPRF gel is obviously increased.

2. The gel network formed by the colloid particles endows the double-network hydrogel with the characteristics of injectability and plasticity:

the nano particles with the size from nanometer to submicron scale can form colloidal gel with shear thinning and self-repairing mechanical characteristics through close packing and physical interaction among the particles, and the gel process of iPRF requires time window property. This is of great convenience in clinical use. The double-network structure enables the colloidal particle/iPRF composite hydrogel to have more excellent mechanical properties, and expands the application range of the conventional biomaterial based on iPRF and colloidal gel.

3. The colloidal particle/iPRF double-network hydrogel has bioactivity and can induce tissue regeneration:

compared with the hydrogel or the porous bracket made of the traditional block material, the double-network hydrogel has better protein drug loading and slow release effects, because the colloidal gel biomaterial has a high specific surface area due to the particle size of the particles with the micro-nano scale, and a large number of charged groups are arranged on the surfaces of the colloidal particles, the colloidal gel biomaterial is easy to form electrostatic adsorption with the protein drugs, and meanwhile, the double-network hydrogel constructed by the blood extract iPRF and the colloidal particles can be directly extracted from the venous blood of a patient by utilizing the characteristics of various growth factors enriched in the blood, so that the high cost of using exogenous or recombinant protein growth factors and the potential immunological rejection problem of using the exogenous protein drugs can be obviously reduced.

4. The preparation method of the obtained double-network structure, reinforced and toughened hydrogel biomedical material is simple and convenient, is beneficial to clinical popularization, is an ideal injectable and shapeable biomedical material for filling, repairing and regenerating human tissue and organ defects, and can be used as an injectable material in the field of implantation materials for repairing and regenerating tissues such as human tissues/organs, injectable hemostatic gels, blood vessel repair and regeneration promoting materials, bones, cartilages, fats, muscles and the like.

Drawings

FIG. 1 is a schematic diagram of the mechanical properties of the hydrogel in example 8; FIG. 1a is an optical picture of injectability, and FIG. 1b is an optical picture of compressive strain and recovery;

FIG. 2 is a growth factor release curve for hydrogel in example 9, comparative example 8;

FIG. 3 is a histological analysis of vascularization and osteogenesis of the sockets 2 weeks after filling the sockets with hydrogel after extraction of the beagle dog of example 10 and comparative example 9; wherein the circles represent newly formed blood vessels and the asterisks represent newly formed bone;

fig. 4 is an optical picture of a 3D bioprinted hydrogel scaffold of example 10; FIG. 4a represents silica particle-IPRF and FIG. 4b represents gelatin particle-IPRF.

Detailed Description

The present invention is further illustrated by the following specific examples.

Example 1

(1) Different sizes of silicon dioxide and iprf

Silica nanoparticles having a particle size of 100,500nm (purchased to sigma chemical company, ltd) were used. The potential of the particles was-30 mV as measured by a nanometer particle sizer.

(2) Silicon dioxide positive electrochemical process

24g of a commercially available 100,500nm silica was weighed, dispersed in 600mL of ethanol, and ultrasonically dispersed for 15min to prepare a 15mg/mL silica suspension. It was placed in a water bath at 40 ℃ and 3mL of silane coupling agent APTMS was added over 20 minutes using a syringe pump, with a stirring speed of 1000 rpm. Followed by reaction at 40 ℃ for 5 h. After the reaction is finished, the particles are centrifugally redispersed and cleaned for 6 times by using ethanol (1 time), ethanol-water (1 time) and water (4 times) respectively, so as to prepare the positively charged silicon dioxide nanoparticles. The potential of the particles was 30mV as measured by a nanometer particle sizer.

(3) Extraction of iPRF

Rabbit blood was drawn through an anticoagulant-free centrifuge tube and iPRF was prepared by centrifugation. The preparation process comprises the following steps: 1) fixing a New Zealand rabbit by using a rabbit fixing frame; 2) removing rabbit hair around middle ear artery of rabbit, and kneading to fully expand middle ear artery; 3) blood is collected from the artery in the ear by 7ml by using a blood taking needle and a centrifuge tube; 4) the mixture was centrifuged in a centrifuge (300g, 3 minutes), and 1ml of the supernatant was collected to obtain iPRF.

(4) Repeatedly blowing and beating the iPRF 1mL obtained by centrifugation with 0.1g of positive charge silicon dioxide particles and 0.1g of negative charge silicon dioxide particles with the same size for 10 times by a luer adapter injector to obtain double-network hydrogel;

(5) the storage and loss moduli of the above-described double-network hydrogels (table 1) were obtained using a time-sweep mode of a rotational rheometer, with a frequency of 1Hz and a strain of 0.5%. The compressive strain at break was obtained by testing the mechanical compression, where the compression rate was 5mm/min and the compressed sample was a cylinder with a diameter of 12mm and a height of 8 mm. It was found that the smaller the particle size, the stronger the composite hydrogel obtained at the same mass fraction.

TABLE 1

	100nm silica particles	500nm silica particles
			Storage modulus	6.3kPa	4.2kPa
Loss modulus	0.54kPa	0.48kPa
			Strain at compression break	81％	79％

Example 2

(1) Different proportions of silica

(2) Silicon dioxide positive electrochemical process

(3) Extraction of iPRF

(4) Repeatedly blowing and beating 1mL of iPRF obtained by centrifugation and 0.16g of positive charge silica particles and 0.04g of negative charge silica particles with the same size for 10 times by using a luer adapter injector to obtain double-network hydrogel;

(5) the storage and loss moduli of the above-described double-network hydrogels (table 2) were obtained using a time-sweep mode of a rotational rheometer, with a frequency of 1Hz and a strain of 0.5%. The compressive strain at break was obtained by testing the mechanical compression, where the compression rate was 5mm/min and the compressed sample was a cylinder with a diameter of 12mm and a height of 8 mm. Compared with Table 2, the strength of the composite hydrogel is reduced after the same mass fraction of positively and negatively charged silica is from 1:1 to 4: 1.

TABLE 2

	100nm silica particles	500nm silica particles
			Storage modulus	4.3kPa	3.4kPa
Loss modulus	0.51kPa	0.42kPa
			Strain at compression break	75％	71％

Example 3

(1) Different sizes of silicon dioxide and iprf

(2) Silicon dioxide positive electrochemical process

(3) Extraction of iPRF

(4) Repeatedly blowing and beating the iPRF 1mL obtained by centrifugation with 0.2g of positive charge silicon dioxide particles and 0.2g of negative charge silicon dioxide particles with the same size for 10 times by a luer adapter injector to obtain double-network hydrogel;

(5) the storage and loss moduli of the above-described double-network hydrogels (table 3) were obtained using a time-sweep mode of a rotational rheometer, with a frequency of 1Hz and a strain of 0.5%. The compressive strain at break was obtained by testing the mechanical compression, where the compression rate was 5mm/min and the compressed sample was a cylinder with a diameter of 12mm and a height of 8 mm.

TABLE 3

	100nm silica particles	500nm silica particles
			Storage modulus	319.3kPa	294kPa
Loss modulus	26.2kPa	23.4kPa
			Strain at compression break	43％	39％

Example 4

(1) Different sizes and different concentrations of silicon dioxide and iprf

Using silica nanoparticles purchased with a particle size of 100,500nm (purchased to Sigma chemical technology Co., Ltd.), the potential of the particles obtained by the nano-particle sizer test was-30 mV.

(2) Silicon dioxide positive electrochemical process

24g of a commercially available 100,500nm silica was weighed, dispersed in 600mL of ethanol, and ultrasonically dispersed for 15min to prepare a 15mg/mL silica suspension. It was placed in a water bath at 40 ℃ and 3mL of silane coupling agent APTMS was added over 20 minutes using a syringe pump, with a stirring speed of 1000 rpm. Followed by reaction at 40 ℃ for 5 h. After the reaction, the mixture was centrifuged and subdivided for 6 times using ethanol (1 time), ethanol-water (1 time) and water (4 times), respectivelyDispersing and cleaning particles to prepare SiO with positive charge₂The surface charge of the particles is 30mV

(3) Extraction of iPRF

Rabbit blood was drawn through an anticoagulant-free centrifuge tube and iPRF was prepared by centrifugation. The preparation process comprises the following steps: 1) fixing a New Zealand rabbit by using a rabbit fixing frame; 2) the rabbit hair around the middle ear artery was removed and kneaded to fully dilate the middle ear artery. 3) Blood is collected from the artery in the ear by 7ml by using a blood taking needle and a centrifuge tube; 4) 0.2g of positively charged silica particle powder and 0.2g of negatively charged silica particle powder were added to a blood collection tube containing whole blood, and centrifuged in a centrifuge (300g, 3 minutes), and the double-network hydrogel was obtained as the upper layer.

(4) The storage modulus and loss modulus of the above-described double-network hydrogel were obtained using a time-sweep mode of a rotational rheometer, with a frequency of 1HZ and a strain of 0.5%. The compressive strain at break was obtained by testing the mechanical compression, where the compression rate was 5mm/min and the compressed sample was a cylinder with a diameter of 12mm and a height of 8 mm. (Table 4)

TABLE 4

	100nm silica particles	500nm silica particles
			Storage modulus	352.4kPa	311.2kPa
Loss modulus	42.2kPa	37.1kPa
			Strain at compression break	51％	43％

Comparative example 1

Rabbit blood was drawn through an anticoagulant-free centrifuge tube and iPRF was prepared by centrifugation. The preparation process is as follows. 1) New Zealand rabbits were fixed using a rabbit mount. 2) The rabbit hair around the middle ear artery was removed and kneaded to fully dilate the middle ear artery. 3) 7ml of blood was collected from the middle ear artery using a blood collection needle and a centrifuge tube. 4) The mixture was centrifuged in a centrifuge (300g, 3 minutes), and 1ml of the supernatant was collected to give iPRF, which was then left for 30min to give an iPRF gel.

The storage and loss moduli of the iPRF were obtained using a time-sweep mode of a rotational rheometer (table 5) with a frequency of 1HZ and a strain of 0.5%. The compressive strain at break was obtained by testing the mechanical compression, where the compression rate was 5mm/min and the compressed sample was a cylinder with a diameter of 12mm and a height of 8 mm.

TABLE 5

	iPRF
		Storage modulus	0.11kPa
Loss modulus	0.02kPa
		Strain at compression break	92％

Comparative example 2

Equal mass of 0.1g or 0.2g of positive and negative charge silicon dioxide powder is mixed with 1mL of deionized water solution to obtain silicon dioxide colloidal gel.

The storage and loss moduli of the gelatin colloidal gels were obtained using a time-sweep mode of a rotational rheometer (table 6) with a frequency of 1HZ and a strain of 0.5%. The compressive strain at break was obtained by testing the mechanical compression, where the compression rate was 5mm/min and the compressed sample was a cylinder with a diameter of 12mm and a height of 8 mm.

TABLE 6

The strength of the silica composite hydrogel in example 1 was significantly higher than the pure iPRF and gelatin nanoparticle hydrogels.

Example 5 Chitosan particle iprf

(1) Preparation of chitosan particles

1g of chitosan powder was weighed and dissolved in 100mL of 1% acetic acid solution to obtain a chitosan solution. Then 20mg of sodium tripolyphosphate powder was weighed and dissolved in 5mL of deionized water to obtain a 4mg/mL sodium tripolyphosphate solution. Stirring the chitosan solution at a high speed of 1000rpm, and dropwise adding 5mL of sodium tripolyphosphate solution to react for 10min to obtain chitosan nanoparticles with positive and negative charge groups locally. The particle size was 200nm and the potential of the particles was-20 mV, as measured by a nanometer particle sizer.

(2) Extraction of iPRF

(3) Repeatedly blowing and beating the iPRF 1mL obtained by centrifugation and 0.1g of chitosan particles for 10 times by a luer adapter injector to obtain the double-network hydrogel;

(4) the storage modulus and loss modulus of the above-described double-network hydrogel were obtained using a time-sweep mode of a rotational rheometer, with a frequency of 1HZ and a strain of 0.5%. The compressive strain at break was obtained by testing the mechanical compression, where the compression rate was 5mm/min and the compressed sample was a cylinder with a diameter of 12mm and a height of 8 mm. (Table 8)

TABLE 8

Storage modulus	7.9kPa
		Loss modulus	1.2kPa
Strain at compression break	78％

Example 6 different concentrations of Chitosan particles Ipff

1) Preparation of chitosan particles

(2) Extraction of iPRF

(4) Repeatedly blowing and beating the iPRF 1mL obtained by centrifugation and 0.2g of chitosan particles for 10 times by a luer adapter injector to obtain the double-network hydrogel;

(5) the storage modulus and loss modulus of the above-described double-network hydrogel were obtained using a time-sweep mode of a rotational rheometer, with a frequency of 1HZ and a strain of 0.5%. The compressive strain at break was obtained by testing the mechanical compression, wherein the compression rate was 5mm/min and the compressed sample was a cylinder with a diameter of 12mm and a height of 8mm (Table 9).

TABLE 9

Storage modulus	36.9kPa
		Loss modulus	7.2kPa
Strain at compression break	58％

Example 7 different mixing modes of chitosan particle iprf

(1) Preparation of chitosan particles

(2) Extraction of iPRF

Rabbit blood was drawn through an anticoagulant-free centrifuge tube and iPRF was prepared by centrifugation. The preparation process comprises the following steps: 1) fixing a New Zealand rabbit by using a rabbit fixing frame; 2) removing rabbit hair around middle ear artery of rabbit, and kneading to fully expand middle ear artery; 3) blood is collected from the artery in the ear by 7ml by using a blood taking needle and a centrifuge tube; 4) 0.2g of chitosan particle powder was added to a blood collection tube containing whole blood, and centrifuged in a centrifuge (300g, 3 minutes), and the double-network hydrogel was obtained as the upper layer.

(3) The storage modulus and loss modulus of the above-described double-network hydrogel were obtained using a time-sweep mode of a rotational rheometer, with a frequency of 1HZ and a strain of 0.5%. The compressive strain at break was obtained by testing the mechanical compression, where the compression rate was 5mm/min and the compressed sample was a cylinder with a diameter of 12mm and a height of 8 mm. (watch 10)

Watch 10

Storage modulus	16.5kPa
		Loss modulus	1.1kPa
Strain at compression break	73％

Comparative example 6 Chitosan nanoparticle

(1) Preparation of chitosan particles

(2) Repeatedly blowing and beating 0.2g of chitosan particles and 1mL of deionized water for 10 times through a luer adapter injector to obtain the double-network hydrogel;

(3) the storage modulus and loss modulus of the above-described double-network hydrogel were obtained using a time-sweep mode of a rotational rheometer, with a frequency of 1HZ and a strain of 0.5%. The compressive strain at break was obtained by testing the mechanical compression, wherein the compression rate was 5mm/min and the compressed sample was a cylinder with a diameter of 12mm and a height of 8mm (Table 11). By comparing the chitosan colloidal gel and the chitosan colloidal particle-iPRF, the storage modulus of the colloidal particle-iPRF group is 25 times higher than that of the colloidal particle and the compressive fracture strain is 1 time higher under the same chitosan mass fraction.

TABLE 11

Storage modulus	1.5kPa
		Loss modulus	0.2kPa
Strain at compression break	28％

Example 8

(1) Preparation of gelatin granules

Dissolving 5g of gelatin in 100mL of deionized water solution, keeping heating to 40 ℃ to obtain clear and transparent gelatin aqueous solution, dropwise adding hydrochloric acid to adjust the pH value of the solution to 2.5, dropwise adding 300mL of acetone solution into the gelatin aqueous solution, keeping heating to 40 ℃ and continuously stirring (1000rpm), dropwise adding for a total time of 20min, adding 74 mu L of cross-linking agent glutaraldehyde (25 wt% aqueous solution) into the nanoparticle suspension, and carrying out cross-linking for 12hrs, after the reaction is finished, adding 100mM glycine into the mixture, and terminating the end group of the glutaraldehyde which is not completely reacted. The nanoparticle suspension was repeatedly centrifuged and resuspended in deionized water. Obtaining the gelatin nano-particles with positive and negative charge groups locally. The potential of the particles was +8.5mV, measured by a nanometer particle sizer, where the size of the particles was 200 nm.

(2) Extraction of iPRF

(3) Repeatedly blowing and beating the centrifuged iPRF 1mL and 0.1g of gelatin nanoparticles for 10 times by a luer adapter injector to obtain gelatin particle-iPRF hydrogel, wherein the hydrogel has shear thinning and injectable mechanical properties, and FIG. 1a represents an injectable schematic diagram of the hydrogel; after 15 minutes, the composite hydrogel is formed along with the complete solidification of the iPRF network, and the hydrogel has stable mechanical properties at the moment, wherein fig. 1b represents that the composite hydrogel can be completely restored to an initial state after being subjected to compression deformation by using an iron rod, and shows the stable mechanical properties of the composite hydrogel.

(4) The storage and loss moduli of the above-described double-network hydrogels (table 12) were obtained using a time-sweep mode of a rotational rheometer, with a frequency of 1Hz and a strain of 0.5%. The compressive strain at break was obtained by testing the mechanical compression, where the compression rate was 5mm/min and the compressed sample was a cylinder with a diameter of 12mm and a height of 8 mm.

TABLE 12

	200nm gelatin particles
		Storage modulus	11.4kPa
Loss modulus	0.84kPa
		Strain at compression break	65.5％

Comparative example 7

(1) Preparation of gelatin granules

Dissolving 5g of gelatin in 100mL of deionized water solution, keeping heating to 40 ℃ to obtain clear and transparent gelatin aqueous solution, dropwise adding hydrochloric acid to adjust the pH value of the solution to 2.5, respectively dropwise adding 300mL of acetone solution into the gelatin aqueous solution, keeping heating to 40 ℃ and continuously stirring (1000rpm), dropwise adding for a total time of 20min, adding 74 mu L of cross-linking agent glutaraldehyde (25 wt% aqueous solution) into the nanoparticle suspension, carrying out cross-linking for 12hrs, adding 100mM glycine into the mixture after the reaction is finished, and terminating the end group of the glutaraldehyde which is not completely reacted. The nanoparticle suspension was repeatedly centrifuged and resuspended in deionized water. Obtaining the gelatin nano-particles with positive and negative charge groups locally. The potential of the particles was +8.5mV, measured by a nanometer particle sizer, where the size of the particles was 200 nm. .

(2) Repeatedly blowing and beating 0.2g of gelatin particles and 1mL of deionized water for 10 times through a luer adapter injector to obtain the double-network hydrogel;

(3) the storage modulus and loss modulus of the above-described double-network hydrogel were obtained using a time-sweep mode of a rotational rheometer, with a frequency of 1HZ and a strain of 0.5%. The compressive strain at break was obtained by testing the mechanical compression, where the compression rate was 5mm/min and the compressed sample was a cylinder with a diameter of 12mm and a height of 8 mm. (Table 13) while observing the pure gelatin particles, the gel strength of the IPRF gel is much lower than that of the gelatin particle-IPRF composite gel.

Watch 13

	200nm gelatin particles
		Storage modulus	0.9kPa
Loss modulus	0.24kPa
		Strain at compression break	17％

Example 9 degradation experiments

The degradation experiments were carried out on pure iprf gel in PBS solution using a colloidal gel of positively and negatively charged silica of 100nm in example 1, iprf gel.

The double-network hydrogels of examples 1-6 were used to determine their growth factor release profiles. Taking the 100nm positively and negatively charged silica nanoparticles of example 1 and the gelatin nanoparticles of example 8 as examples, two sets of samples were soaked in phosphate buffered saline (PBS, pH 7.4) and kept at 37 ℃ ambient temperature on a shaker (30 rpm) to simulate in vivo dynamic environment. 1ml of PBS supernatant was taken up at 1d, 3d, 7d, 14d and 21d, respectively, and then an equal amount of fresh PBS was added. TGF- β was examined at each time point using an ELISA kit according to the manufacturer's instructions, all experiments with VEGF were performed in triplicate, with three samples per group. As shown in fig. 1, VEGF and TGF- β in the hydrogels of iPRF-gelatin particles and iPRF-silica particles were released at a relatively constant concentration as measured by ELISA, and this uniform release was still detectable at day 15.

Comparative example 8 growth factor Release

The iPRF gel of comparative example 1 was used to determine its growth factor release profile. Both sets of samples were soaked in phosphate buffered saline (PBS, pH 7.4) and kept at 37 ℃ ambient temperature on a shaker (30 rpm) to simulate an in vivo dynamic environment. 1ml of PBS supernatant was taken up at 1d, 3d, 7d, 14d and 21d, respectively, and then an equal amount of fresh PBS was added. TGF- β was examined at each time point using an ELISA kit according to the manufacturer's instructions, all experiments with VEGF were performed in triplicate, with three samples per group. And (3) releasing a large amount of factors from the iPRF gel in the third day, releasing the rest factors from the fibrin network of the iPRF in a concentration decreasing mode, and finally releasing the active factors at the 14 th day, wherein the active factors are released quickly in a short term and cannot be released slowly for a long time compared with the colloidal particle-iPRF composite gel.

Example 10

Beagle dogs were anesthetized with 3 wv% sodium pentobarbital (1 ml/kg). The mandibular fourth premolar was extracted and 2 sockets were obtained from each side. The left socket was filled with the double-network hydrogel of example 8. Histological analysis was performed at the site of tooth extraction at 2 weeks.

As shown in the results of fig. 3, the double-network hydrogel prepared in example 8 has a significantly better bone effect than the gel of the gelatin particles alone in comparative example 7. The circles in FIG. 2a represent new blood vessels, the stars in FIG. 2b represent newly formed bone regions, and at 2w, it can be seen that the gelatin particle-IPRF hydrogel group has more blood vessels and new bone formation. The above experimental results demonstrate that hydrogels of IPRF-GNPs can promote angiogenesis and osteogenesis in localized areas.

Comparative example 9

Beagle dogs were used. The mandibular fourth premolar was extracted and 2 sockets were obtained from each side. The right lateral socket was filled with gelatin particle gel of comparative example 7. Histological analysis was performed at the site of tooth extraction at 2 weeks. As shown in the results of fig. 3, the gelatin particle gel did not have angiogenisis and osteogenesis effects at 2 w.

Example 11

Using the two-network hydrogel obtained in examples 1 to 8, the silica particle-IPRF in example 1 and the gelatin particle-IPRF in example 8 were used as examples, the polymerization of the fibrinogen network was stopped by adding 5mM sodium citrate, the three-dimensional bio-printer was used to print under the condition that the hydrogel had shear thinning (i.e., was printable) to obtain a customized scaffold, and the fibrinogen was polymerized to form the fibrin network by soaking in 10mM calcium chloride solution to obtain a reinforced and toughened two-network hydrogel scaffold, wherein the structure of the gelatin particle-IPRF particle-iPRF scaffold is shown in FIG. 4a, and the structure of the silica particle-IPRF scaffold is shown in FIG. 4 b.

Claims

1. The double-network-structure composite hydrogel based on the colloidal particles and the iPRF is characterized in that the composite hydrogel is formed by injecting platelet-enriched fibrin iPRF based on the colloidal particles and a blood extract and has a double-network microstructure; the gel particles form a first heavy colloidal gel network under the physical action, the injectable platelet-enriched fibrin iPRF forms a second heavy fibrin hydrogel network, and the two heavy gel networks are physically crosslinked by electrostatic action, hydrogen bond action and hydrophobic action to form hydrogel with a double-network structure;

the curing time of the colloidal particle-iPRF double-network structure composite hydrogel is 10-2000 seconds, wherein the injectable and shapeable time window is less than 2000 seconds; the particle size of the colloidal particles is 10 nm-2 mu m;

the injectable platelet-rich fibrin iPRF is a blood extract from an autologous, allogeneic or xenogeneic source;

the colloidal particles are assembled by electrostatic interaction, and are specifically combined by any one of the following modes:

2) the nano-particle is prepared from one or more of montmorillonite, chitosan, alginic acid, hyaluronic acid, fibroin and gelatin nano-particles which are locally provided with positive charge and negative charge groups.

2. The composite hydrogel according to claim 1, wherein the surface charge of the colloidal particles is-40 to 40 mV.

3. The composite hydrogel according to claim 1, wherein the colloidal particles have a particle size of 10nm to 500 nm.

4. The composite hydrogel according to claim 1, wherein the double-network structure composite hydrogel is used as a carrier of drugs, bioactive proteins and living cells, and is applied to fillers of bone tissue, cartilage tissue, muscle and vascular repair.

5. A method for preparing the colloidal particle-iPRF based double network structure composite hydrogel according to any one of claims 1 to 4, wherein the method comprises the following steps:

Adding colloidal particles into a fresh blood sample, placing the sample into a centrifuge for centrifugation for 1-10 min at a rotation speed of 50-1000 g, and solidifying a top brown yellow colloid after centrifugal layering to obtain the colloidal particle-iPRF double-network structure composite hydrogel;

wherein the volume fraction phi of the colloidal particles is phi = 0.05-1.

6. The method for preparing the composite hydrogel according to claim 5, wherein the colloidal particles are one or a combination of a dispersion of the colloidal particles in an aqueous solution and a lyophilized powder of the colloidal particles.

7. The application of the colloidal particle-iPRF-based double-network-structure composite hydrogel of claim 1 in preparation of three-dimensional cell culture scaffolds or 3D bioprinting ink materials added to cells for 3D bioprinting of tissues or organs, wherein 1-100 mmol of anticoagulant is added to the double-network-structure composite hydrogel before printing, and the double-network-structure composite hydrogel is soaked in Ca-containing solution after printing²⁺In aqueous solution, a scaffold with a stable structure is obtained.