A collagen/hydroxyapatite composite scaffold, and process for the production thereof

Technical field [0001] The invention relates to a method for producing collagen / hydroxyapatite (HA) composite scaffolding as well as collagen / HA composite scaffolds obtainable by the method of the invention. Such scaffolds can be used in bone regeneration and tissue engineering applications.

Background of the Invention [0002] Bone grafts are the second most commonly transplanted materials after the blood transfusion worldwide. In addition, the estimated value of the global bone graft material market is around 650 million euros per year. Every year, up to 4 million bone replacement procedures are performed around the world that require bone grafting or scaffolding. The most common clinical therapy is autograft, in which the bone is taken from the patient's own body and re-implanted. However, the amount of bone that can be removed from a given donor site is limited, and re-implantation requires additional invasive surgery. Another option is the use of an allograft, in which the bone is taken from an organ donor. The problems associated with this approach arise from the origin of the bone from a separate donor. Such material is associated with an increased risk of transmitting infectious diseases. In addition, there are fewer growth factors in such a donor bone because it contains no viable cells. These growth factors support new bone growth. An ideal implanted scaffold that would stimulate bone formation while facilitating load bearing would reduce the need for allografts or autografts. Currently, however, these traditional approaches account for over 90% of all bone grafting procedures. The reason for this deficiency, in addition to the problems described above, is that a vascularized, mechanically appropriate, osteoconductive scaffold that could be used to produce bone in vitro or to induce complete osteogenesis in vivo still remains to be developed. Such a product will have significant commercial potential. [0003] Various tests have been carried out using many synthetic materials for the production of useful bone graft scaffolds. Examples include polystyrene, titanium, polylactic acid (PLLA), polyglycolic acid (PGA) and lactic acid copolymer with glycolic acid (PLGA). However, all these materials have associated problems and inconveniences, including the risk of infection and the difficulty of providing adequate resorption to promote vascularization and ingrowth of new bone. Substrates based on biological materials such as collagen, gelatin, chitosan, agarose and glycosaminoglycans (GAG) are also used. However, these materials do not have sufficient mechanical properties to be able to bear the load after implantation. WO2006 095154 describes calcium phosphate and collagen composites that are prepared using mechanical mixing or by precipitation in situ of a HA-like phase.

Summary of the Invention [0004] According to the invention, there is provided a process for producing a collagen / hydroxyapatite (HA) composite scaffold according to claims 1 to 9.

[0005] Typically, the acid solution has a molar concentration of at least 0.05 M. In a preferred embodiment of the invention, the HA to collagen suspension ratio is greater than 1:10 (w / w), wherein the molar concentration of the acid solution is greater than 0.05 M. Usually the ratio of HA to collagen in suspension is at least 2:10 (w / w), 3:10 (w / w), 4:10 (w / w), 5: 10 (w / w). In one preferred embodiment of the invention, the ratio of HA to collagen is from 1:10 (w / w) to 50:10 (w / w), respectively from 5:10 (w / w) to 30:10 ( w.). The molar concentration of the acid solution, respectively, is at least 0.06 M, 0.07 M, 0.08 M, 0.09 M, 0.10 M, 0.20 M, 0.30 M, 0.40 M or

0.50 M. Ideally, the molar concentration of the acid solution is between 0.4 M and 0.6 M.

[0006] In one preferred embodiment of the invention, the HA to collagen suspension ratio is at least 5:10 (w / w), wherein the molar concentration of the acid solution is at least 0.10 M. Typically, the molar concentration of the acid solution is at least 0.50 M.

[0007] In a preferred embodiment of the invention, the ratio of HA to collagen in suspension is at least 6:10 (w / w), 7:10 (w / w), 8:10 (w / w), 9:10 (w / w) or 1: 1 (w / w). In one embodiment of the invention, the HA to collagen suspension ratio is greater than 1: 1 (w / w). Generally, when such HA levels are used in suspension, the molar concentration of the acid solution will be less than 0.5 M.

[0008] In a preferred embodiment, the amount of collagen in the suspension may range between 4.0 g / L and 6.0 g / L.

[0009] Typically, the acidic solution includes an acetic acid solution. However, other organic acids can be used to form the acid solution.

[0010] Correspondingly homogeneous collagen / HA suspension is formed under appropriate conditions minimizing collagen gelatinization. One way to ensure minimal collagen gelatinization during the production of a homogeneous suspension is to keep the suspension at a sufficiently low temperature, generally between 1 ° C and 5 ° C, about 4 ° C, respectively.

[0011] Lyophilization is carried out at a constant cooling rate. This means that the cooling rate during freeze drying does not change by more than +/- 10% of the target cooling rate, i.e. if the desired cooling rate is 1.0 ° C / min and the actual cooling rate is between 0.9 ° C / min and 1.1 ° C / min, it is still considered a constant cooling rate. Typically, the constant cooling rate is from 0.1 ° C / min to 10 ° C / min. Preferably, lyophilization is carried out at a constant cooling rate between 0.5 ° C / min and 1.5 ° C / min. More preferably, lyophilization is carried out at a constant cooling rate between 0.8 ° C / min and 1.1 ° C / min. Typically, lyophilization is carried out at a constant cooling rate of about 0.9 ° C / min. The temperature in the freeze drying chamber at the beginning of the freeze drying process (i.e. when the suspension is placed in the drying chamber) is usually higher than 0 ° C, preferably it is near ambient temperature.

[0012] The desired final freezing temperature is between -10 ° C and -70 ° C. Accordingly, the desired final freezing temperature is between -30 ° C and -50 ° C. Usually the desired final freezing temperature is between -35 ° C and -45 ° C, ideally around -40 ° C.

[0013] The freeze-drying process comprises a drying step which is carried out after reaching the final freezing temperature. This step involves heating the lyophilization chamber to a sublimation temperature (usually around 0 ° C), preferably at a constant heating rate. This process usually involves the final stage of sublimation, in which the ice phase in the scaffolding formed sublimates under vacuum for a suitable time.

[0014] In another embodiment of the invention, the freeze-drying process comprises a soaking step. Typically, this step involves increasing the temperature in the freeze drying chamber after reaching the final freezing temperature and usually maintaining the elevated temperature for some time before starting the drying step. For example, when the final freezing temperature is -20 ° C, the soaking step can be carried out by raising the temperature to -10 ° C and maintaining at this temperature for a time sufficient to allow existing ice crystals to grow before the scaffold is finally dried. The soaking time can be varied depending on the desired pore properties, however, soaking time between 15 minutes and 120 hours is preferred.

[0015] Generally, the HA used in the present invention is in the form of a powder. Accordingly, the HA powder is selected from the group consisting of: sintered HA powder; and sintered HA powder. Examples of suitable sintered and sintered HA powders suitable for the present invention are known to the skilled person and are described below.

[0016] Typically, the HA powder has a particle size in the range between 10 nm and 100 Pm.

[0017] Accordingly, the collagen used in the present invention includes collagen fibers. Preferably the collagen fibers include microfibrillar collagen, preferably microfibrillar collagen from bovine tendons.

[0018] Suitably, the collagen suspension is centrifugally mixed, with HA being added to the vortex in the suspension under centrifugal mixing.

[0019] Typically, HA is added in portions. Correspondingly, portions are added to the collagen suspension at intervals of between 30 and 240 minutes. Preferably HA is added to the collagen suspension in 2 to 5 portions.

[0020] In one embodiment of the invention, the composite scaffold is crosslinked. Usually the composite scaffold is crosslinked in a way selected from the group consisting of: dehydrothermal crosslinking and chemical crosslinking. Suitable chemical crosslinkers and methods are well known to those skilled in the art and include 1-ethyl-3- [3-dimethylaminopropyl] carbodiimide hydrochloride (EDAC). When dehydrothermal crosslinking is used, the crosslinking temperature is between 105 ° C and 180 ° C. Preferably, the crosslinking process is carried out for at least 24 hours, 48 ​​hours, 72 hours, 96 hours or 120 hours. When EDAC crosslinking is used, the molar concentration of the EDAC solution is 6 mmol per gram of collagen / HA composite.

[0021] The invention also relates to a collagen / hydroxyapatite (HA) composite scaffold which can be obtained by the process of the invention.

[0022] The invention also relates to a collagen / hydroxyapatite (HA) composite scaffold according to claims 10 to 13.

[0023] Preferably, the composite scaffold according to the invention has a porosity of at least 99% (v / v), 99.1% (v / v), 99.2% (v / v), 99.3 % (v / v). Accordingly, the composite scaffold according to the invention has a porosity of 98 to 99.5% (v / v), and more preferably of 98.5 to 99.5% (v / v). The method for determining% porosity is described below.

[0024] Suitably, the composite scaffold according to the invention has a compressive stiffness of at least 0.5 KPa, 0.6 KPa. Accordingly, the composite scaffold according to the invention has a compressive stiffness in the range from 1 to 5 KPa, preferably from 1 to 4 KPa. EDAC crosslinked composite scaffoldings have a compressive stiffness of at least 1 kPa, 1.5 kPa, 2 kPa, 2.5 kPa, 3 kPa, 3.5 kPa, 4 kPa. The method of determining compression stiffness is described below.

[0025] Typically, the ratio of HA to collagen in the composite scaffold is from 1:10 to 50:10 (w / w), preferably at least 2:10 (w / w), 3:10 (w / w) w / w), 4: 10 (w / w), 5:10 (w / w), 6:10 (w / w), 7:10 (w / w), 8: 10 (w / w), 9:10 (w / w) or 1: 1 (w / w). In a particularly preferred embodiment of the invention, the ratio of HA to collagen in the composite scaffold is from 5:10 to 30:10 (w / w).

[0026] The in vitro biological activity of the composite scaffold of the invention can be characterized by monitoring osteoblast activity of MC3T3 in the scaffold after 1 day (to monitor the degree of initial cell attachment), 7 days, 21 days and 28 days of incubation (to monitor cell proliferation). In one embodiment of the invention, the composite scaffold of the invention is characterized by a level of osteoblast proliferation MC3T3 in the scaffold after 7 days of incubation greater than the initial number of cells seeded on the scaffold. It is usually at least 1x10⁶ cells per 500 mm³ scaffolding volume. In a preferred embodiment of the invention, the composite scaffold is characterized by a difference between the osteoblast proliferation level MC3T3 in the scaffold after 28 days of incubation and the level after 7 days of at least 0.5 x 10⁶ cells, respectively from 0.5 x 10⁶ up to 1.5 x 10⁶ cells. The method of determining the level of osteoblast proliferation of MC3T3 is as follows: 2 x 10 cylindrical scaffold samples with a diameter of 12.7 mm are sown⁶ MC3T3 cells.

After 7 days of incubation, the number of cells present on the scaffold is monitored using the Hoechst 33258 DNA test. It gives an assessment of the initial cell attachment. After 14, 21 and 28 days of incubation, the number of cells present on the scaffold is monitored using the Hoechst 33258 DNA test. The change in the number of cells present on the scaffold over time (number of cells on day 28 minus number of cells on day 7) is used to assess cell proliferation .

[0027] In one embodiment, the composite scaffold according to the invention has a pressure conductivity through the scaffold of at least 1 x 10¹⁰ m⁴/ Ns., From 6 x 10 respectively^-10 m⁴/ Ns up to 1.4 x 10^-9 m⁴/ Ns, preferably from 8 x 10^-10m⁴/ Ns up to 1.2 x 10^-9 m⁴/ Ns. Typically, the conductivity of pressure under pressure through the scaffolding is at least 10 x 10^-10 m⁴/ Ns.

[0028] Ideally, the composite scaffold according to the invention has a high degree of interconnection between the pores. Preferably the scaffold has an even distribution of pores. Usually, the scaffolding has a homogeneous pore size. In one embodiment, the scaffold is manufactured in the form of a sheet. Typically, the sheet has an average thickness of at least 1 mm, 2 mm, 3 mm, 4 mm, 5 mm, 6 mm, 7 mm, 8 mm, 9 mm or 10 mm.

[0029] The invention also relates to a composite scaffold according to the invention or obtainable by the method according to the invention on which cells have been seeded. These cells are usually stem cells that are undifferentiated, partially differentiated or completely differentiated. In one embodiment, the cells are selected from the group consisting of: osteoblasts and mesenchymal stem cells.

[0030] The invention also relates to an osteoconductive bone implant comprising a composite scaffold according to the invention.

[0031] The invention also relates to a tissue engineering implant comprising a composite scaffold according to the invention. Thus, the scaffold according to the invention may form the basis on which tissue can be constructed. Various forms of tissue are envisaged for this use, including but not limited to cartilage, ligaments, muscles, and organs.

[0032] The invention also relates to a bone substitute for maxillofacial transplants comprising a composite scaffold according to the invention.

[0033] The invention also relates to a bone substitute for dental implants comprising a composite scaffold according to the invention.

[0034] The invention also relates to an cartilage defect repair implant comprising the composite scaffold according to the invention.

[0035] The invention also relates to an bone-cartilage defect repair implant comprising a composite scaffold according to the invention.

[0036] The proposed invention is made of two basic bone components, namely the mineral phase, hydroxyapatite (HA) and the organic phase, collagen. As such, it is a more inert substrate than any of the previously described materials that stimulate bone formation. Furthermore, by combining the high mechanical stiffness of HA with the biocompatibility, biodegradability and pore architecture of the collagen scaffold produced using the specific method of the invention, a product has been developed that meets all criteria required for use as an osteoconductive scaffold, including excellent compressive stiffness (for ease of manipulation and in vivo loading) and a high degree of porosity, interconnection between pores and permeability.

[0037] Hydroxyapatite is a ceramic material. Ceramic materials are inorganic, non-metallic compounds that form ionic and covalent bonds. They are characterized by high mechanical stiffness, very low elasticity and a hard, brittle surface. In living tissue, HA combines with collagen to form basic bone components. As a material, it also has both chemical and crystalline resemblance to bone minerals. However, pure HA constructs are unattractive for many reasons, primarily due to the stiffness, brittle nature and poor resorbability of the material [1]. Therefore, stability and control of the degradation rate of the construct are problematic [2], strongly inhibiting optimal resorption, subsequent tissue ingrowth and restoration of the mechanical integrity of the defect site, all of which are important determinants of successful implantation.

[0038] Unlike HA, the second component of the present invention, collagen, already meets all the biological determinants required for successful implantation. It is a natural polymer present in many tissues throughout the human body, thus showing excellent biocompatibility. As a result, collagen stimulates cell adhesion, proliferation and intercellular substance formation (ECM). Its degradation rate can be controlled in vivo by changing the crosslinking density. Crosslinking bonds are chemical bonds between collagen molecules. They provide collagen mechanical strength and stabilize collagen fibers, preventing long rod-shaped collagen particles from shifting against each other under stress [3]. Cross-linking is also an effective way to control the rate of degradation of collagen scaffolds, since cross-linking bonds must be broken before the scaffold can degrade.

[0039] There are various ways to increase the level of crosslinking in collagen scaffolds. Another important feature is the recent FDA approval and clinical success of collagen-based scaffolds used for skin and nerve regeneration. [4]. The main disadvantage associated with collagen as a scaffold is the lack of natural mechanical strength. Therefore, the present invention combines both collagen and HA to form a three-dimensional composite construct having the advantages of both components and having none of their disadvantages.

[0040] In addition to the actual component materials themselves, the scaffold manufacturing method and the resulting morphology of the construct are important in determining the in vivo success of the scaffold implant. The production of the present invention involves the use of a specialized collagen scaffold manufacturing technique, which typically includes freeze drying / freeze drying.

[0041] The production of porous scaffolds using freeze-drying has traditionally involved rapid freezing or cooling of scaffolding constituent materials mixed together in suspension. This leads to a very irregular pore distribution and a large degree of variability in pore size. In addition, cooling changes the elongation of the pores formed, which leads to unequal pore shapes throughout the material. The freeze-drying production process of the present invention facilitates full control of all major morphological indicators of scaffold life. This is achieved by accurately controlling the temperature and pressure in the freeze drying chamber during both the freezing and drying stages. It has been shown that uncontrolled freezing or drying at any time during the production process leads to heterogeneous distribution of pores, their shapes and sizes, all of which are key determinants of the viability of seed cells. By using a controlled freeze-drying process, it is possible to reproducibly produce porous, repeatable and homogeneous scaffolds based on collagen with a high degree of interconnection between pores, porosity and surface area, all of which are necessary for efficient transport of cell mass within the scaffold and surrounding host tissue and provide space for vascularization and growth of new tissue. Furthermore, the method of the invention allows extensive control of the pore size of the construct, facilitating cell-specific functionality. The present disclosure describes the production of a composite scaffold by using the method of the invention, which results in a porous, three-dimensional scaffold with high porosity, a high degree of interconnection between pores and a homogeneous distribution of HA in the collagen matrix.

Short description of the figures [0042]

Figure 1: Stiffness in compression of the non-crosslinked EDAC scaffold depending on the type of scaffold. (collagen = control collagen scaffold mixed in 0.05 M acetic acid, 10 wt% HA = collagen + 10 wt% HA mixed in 0.05 M acetic acid, 50 wt% HA (L) = collagen + 50% by weight HA mixed in 0.1 M acetic acid, 50% by weight HA = collagen + 50% by weight HA mixed in 0.5 M acetic acid).

Figure 2: Stiffness in compression of the EDAC cross-linked scaffold depending on the type of scaffold. (collagen = EDAC cross-linked collagen scaffold mixed in 0.5 M acetic acid, 50 HA = EDAC cross-linked collagen + 50% by weight HA mixed in 0.5 M acetic acid, 100 HA = EDAC cross-linked collagen + 100% by weight HA mixed in 0.5 M acetic acid, 200 HA = EDAC cross-linked collagen + 200 wt% HA mixed in 0.5 M acetic acid).

Figure 3: Porosity of the scaffold depending on the type of scaffold. (collagen = non-cross-linked EDAC collagen scaffold mixed in 0.5 M acetic acid, 50 HA = non-cross-linked EDAC collagen scaffold + 50% by weight HA mixed in 0.5 M acetic acid, 100 HA = non-cross-linked EDAC collagen scaffold + 100% by weight HA mixed in 0.5 M acetic acid, 200 HA = EDAC uncrosslinked collagen scaffold + 200 wt% HA mixed in 0.5 M acetic acid).

Figure 4: In vitro biological activity depending on the type of scaffold (t = 7, 14, 21, 28 days). Initial seeding density of 2 million cells was used for all scaffolds. (collagen = EDAC cross-linked collagen scaffold mixed in 0.5 M acetic acid, 50 HA = EDAC cross-linked collagen + 50% by weight HA mixed in 0.5 M acetic acid, 100 HA = EDAC cross-linked collagen + 100% by weight HA mixed in 0.5 M acetic acid, 200 HA = EDAC cross-linked collagen + 200 wt. HA mixed in 0.5 M acetic acid).

Figure 5: In vitro biological activity depending on the type of scaffold (t = 28 days - 7 days) New scaffold biological activity between day 7 and day 28 (collagen = EDAC cross-linked collagen scaffold scaffold mixed in 0.5 M acetic acid, 50 HA = EDAC cross-linked collagen + 50% by weight HA mixed in 0.5 M acetic acid, 100 HA = EDAC cross-linked collagen + 100% by weight HA mixed in 0.5 M acetic acid, 200 HA = EDAC cross-linked collagen + 200% by weight . HA mixed in 0.5 M acetic acid).

Figure 6: MicroCT scan of the composite scaffold according to the invention.

Figure 7: Cross-section of the scaffolding of Figure 6.

Figure 8: SEM image of a 50% wt scaffold HA highlighting the homogeneous and interconnected pore structure.

Figure 9: Distribution of mineral particles from the area of ​​interest defined in the figure

8.

Figure 10 and 11: SEM images of a composite scaffold with 50 wt. HA of uncrosslinked EDAC magnified 10 times and 100 times, respectively.

Figure 12: Scaffold permeability depending on the scaffold type. (collagen = EDAC cross-linked collagen scaffold mixed in 0.5 M acetic acid, 50 HA = EDAC cross-linked collagen + 50% by weight HA mixed in 0.5 M acetic acid, 100 HA = EDAC cross-linked collagen + 100% by weight HA mixed in 0.5 M acetic acid, 200 HA = EDAC cross-linked collagen + 200 wt% HA mixed in 0.5 M acetic acid).

Figure 13: X-ray image of the rat skull vault bone with an accompanying diagram showing the location of the defect and the location of the sections taken for CT scan images.

Figure 14: MicroCT scan images of sections of bone vaulting rat skull with a void defect.

Figures 15 and 16: MicroCT scan images of sections of the vaulting bone of a rat skull with a defect filled with a 50% wt composite scaffold HA crosslinked

EDAC according to the invention with plated MSC rat cells.

Figure 17 and 18: MicroCT images of sections of bone vaulting of the rat skull with a defect filled with a composite scaffold of 200 wt. HA cross-linked EDAC according to the invention with plated MSC rat cells.

Figures 19 and 20: MicroCT scan images of sections of the vaulting bone of a rat skull with a defect filled with a composite scaffold of 50 wt. HA crosslinked EDAC according to the invention (not shed).

Figures 21 and 22: CT scans of sections of bone vaulting of the rat skull with a defect filled with a composite scaffold of 200 wt. HA crosslinked EDAC according to the invention (not shed).

Detailed description of the invention [0043] Preparation of control collagen scaffolds and scaffolds according to the invention

10 wt. HA was carried out using the protocol described in Form 1, in particular using an initial acetic acid concentration of 0.05 M. As the HA content was increased to 50% by weight, homogeneous mixing of the two main components (collagen and HA) has become more problematic. Increasing the initial acetic acid concentration has been shown to alleviate this problem. The effect of this increased acetic acid concentration was tested using two different increases in acetic acid concentration, in particular 0.1 M and 0.5 M. These forms are described in Examples 2 and 3, respectively. Preparation of EDAC cross-linked collagen scaffold scaffolds and 50% composite scaffolds wt., 100 wt. and 200 wt. HA is described in Example 4.

Example 1 [0044] 400 ml 0.05 M acetic acid solution (pH = 3.05) was prepared using distilled deionized water (1.16 ml glacial acetic acid was added to 398.84 ml distilled deionized water).

[0045] The WKI250 water cooling system (Lauda, ​​Westbury, NY, USA) was used to cool the glass reaction vessel to a constant temperature of 4 ° C for one hour. This reaction vessel was used to mix the scaffold components while maintaining the suspension at a constant temperature of 4 ° C. This prevented denaturation of collagen fibers due to heat generation during the mixing process. 1.8 g of microfibrillar collagen from bovine tendons (Collagen Matrix Inc, NJ, USA) was added to 320 ml of a 0.05 M acetic acid solution. This suspension was mixed using an IKA Ultra Turrax T18 overhead mixer (IKA Works Inc, Wilmington, NC) at 15,000 rpm for 90 minutes at 4 ° C. 40 ml acetic acid solution was mixed with hydroxyapatite (HA) powder (Biotal, UK), exactly 10% by weight collagen (0.18 g HA). A 10 mL aliquot of this acetic acid / HA solution was added to the collagen / acetic acid suspension in the cooled reaction vessel after 90 minutes. The method of delivering the HA suspension included shaking the suspension vigorously (to ensure a homogeneous suspension of mineral particles) immediately before being injected into the center of the vortex in a mixer using a syringe. A flexible rubber tube was attached to the tip of the syringe to facilitate injection directly into the vortex of the mixer. Then 10 ml aliquots (three in total) were added to the suspension every hour. After adding the last portion of the acetic acid solution from HA, the suspension was stirred for another 60 minutes, resulting in a total mixing time of 330 minutes (five and a half hours).

[0046] After completion of the mixing step, the suspension was transferred to a clean, wide neck beaker and vacuum degassed at a pressure of about 4000 mt for an additional 60 minutes. At this stage, all unwanted air bubbles in the suspension have been removed that could have a detrimental effect on subsequent freeze-drying. Scaffolding was made using a lyophilization (freeze drying) process. A 67.5 ml aliquot of collagen / HA suspension was placed on a walled freeze dryer tray provided by the freeze dryer manufacturer (VirTis Co., Gardiner, NY) and made of 304 grade stainless steel. The sample tray had internal dimensions of 127 mm wide x 127 mm long x 38 mm high. The thickness of the plate underlying the tray is 3 mm. The sample tray was placed in the freeze dryer chamber and placed on the cooling shelf of the freeze dryer at 20 ° C.

[0047] The freeze drying process involved cooling the freeze dryer chamber and cooling shelf with a constant cooling rate (0.9 ° C / min), based on previous studies, to a final freezing temperature (-40 ° C). The main determinant of the ice crystal morphology during the freeze drying process is the final freezing temperature. The shelf and chamber temperature was then kept constant at the final freezing temperature for 60 minutes until the freezing process was completed. The shelf temperature was then raised to 0 ° C in 160 minutes. The ice phase was then sublimated under about 200 mt vacuum at 0 ° C for 17 hours to produce a collagen / HA porous scaffold.

[0048] The porous collagen / HA construct was then placed in a vacuum oven (Fisher IsoTemp 201, Fisher Scientific, Boston, MA) to cross-link collagen in a dehydrothermal crosslinking process. The scaffolds were placed in a vacuum oven at 120 ° C under a vacuum of 50 mt for 24 hours.

Example 2 [0049] 400 ml of a 0.1 M acetic acid solution (pH = 2.9) was prepared using distilled deionized water (2.32 ml glacial acetic acid was added to 397.68 ml distilled deionized water). A WKI250 water cooling system (Lauda, ​​Westbury, NY, USA) was used to cool the glass reaction vessel to a constant temperature of 4 ° C for one hour. 1.8 g of microfibrillar collagen from bovine tendons (Collagen Matrix Inc, NJ, USA) was added to 320 ml of a 0.1 M acetic acid solution. This suspension was mixed using an IKA Ultra Turrax T18 overhead mixer (IKA Works Inc, Wilmington, NC) at 15,000 rpm for 90 minutes at 4 ° C.

[0050] 40 ml acetic acid solution was mixed with hydroxyapatite (HA) powder (Biotal, UK), exactly 50% by weight collagen (0.9 g HA). A 10 mL aliquot of this acetic acid / HA solution was added to the collagen / acetic acid suspension in the cooled reaction vessel after 90 minutes. Then 10 ml aliquots (three in total) were added to the suspension every hour. After adding the last portion of the acetic acid solution from HA, the suspension was stirred for another 60 minutes, resulting in a total mixing time of 330 minutes (five and a half hours). After completion of the mixing step, the suspension was transferred to a clean wide neck beaker and degassed under vacuum at about 4,000 mt for an additional 60 minutes.

[0051] Scaffolding was prepared using a lyophilization (freeze drying) process. A 67.5 ml aliquot of collagen / HA suspension was placed on a wall-freeze-dried sample tray provided by the freeze-dried manufacturer (VirTis Co., Gardiner, NY) and made of 304 grade stainless steel. The sample tray had internal dimensions of 127 mm wide x 127 mm long x 38 mm high. The thickness of the plate underlying the tray is 3 mm. The sample tray was placed in the freeze dryer chamber and placed on the cooling shelf of the freeze dryer at 20 ° C.

[0052] The freeze drying process involved cooling the freeze dryer chamber and cooling shelf with a constant cooling rate (0.9 ° C / min) to a final freezing temperature (40 ° C). The shelf and chamber temperature was then kept constant at the final freezing temperature for 60 minutes. The shelf temperature was then raised to 0 ° C in 160 minutes. The ice phase was then sublimated under about 200 mt vacuum at 0 ° C for 17 hours.

[0053] The porous collagen / HA construct was then placed in a vacuum oven (Fisher IsoTemp 201, Fisher Scientific, Boston, MA) to cross-link collagen in a dehydrothermal crosslinking process. The scaffolds were placed in a vacuum oven at 120 ° C under a vacuum of 50 mt for 24 hours.

Example 3 [0054] 400 ml of a 0.5 M acetic acid solution (pH = 2.55) was prepared with distilled deionized water (11.6 ml glacial acetic acid was added to 388.4 ml distilled deionized water). A WKI250 water cooling system (Lauda, ​​Westbury, NY, USA) was used to cool the glass reaction vessel to a constant temperature of 4 ° C for one hour. 1.8 g of microfibrillar collagen from bovine tendons (Collagen Matrix Inc, NJ, USA) was added to 320 ml of a 0.5 M acetic acid solution. This suspension was mixed using an IKA Ultra Turrax T18 overhead mixer (IKA Works Inc, Wilmington, NC) at 15,000 rpm for 90 minutes at 4 ° C.

[0055] 40 ml acetic acid solution was mixed with hydroxyapatite (HA) powder (Biotal, UK), exactly 50%, 100% and 200% by weight collagen (0.9, 1.8, and 3.6 g HA) . A 10 mL aliquot of this acetic acid / HA solution was added to the collagen / acetic acid suspension in the cooled reaction vessel after 90 minutes. Then 10 ml aliquots (three in total) were added to the suspension every hour. After adding the last portion of the acetic acid solution from HA, the suspension was stirred for another 60 minutes, resulting in a total mixing time of 330 minutes (five and a half hours). After completion of the mixing step, the suspension was transferred to a clean wide neck beaker and degassed under vacuum at about 4,000 mt for an additional 60 minutes.

[0056] Scaffolding was prepared using a lyophilization (freeze drying) process. An aliquot of 67.5 ml collagen / HA suspension was placed on the wall of the lyophilizer tray provided by the lyophilizer manufacturer (VirTis Co., Gardiner, NY) and made of 304 grade stainless steel. The sample tray had internal dimensions of 127 mm wide x 127 mm long x 38 mm high. The thickness of the plate underlying the tray is 3 mm. The sample tray was placed in the freeze dryer chamber and placed on the cooling shelf of the freeze dryer at 20 ° C.

[0057] The freeze drying process involved cooling the freeze dryer chamber and cooling shelf at a constant cooling rate (0.9 ° C / min) to a final freezing temperature (40 ° C). The shelf and chamber temperature was then kept constant at the final freezing temperature for 60 minutes. The shelf temperature was then raised to 0 ° C in 160 minutes. The ice phase was then sublimated under about 200 mt vacuum at 0 ° C for 17 hours to produce a collagen / HA porous scaffold.

[0058] The porous collagen / HA construct was then placed in a vacuum oven (Fisher IsoTemp 201, Fisher Scientific, Boston, MA) to cross-link collagen in a dehydrothermal crosslinking process. The scaffolds were placed in a vacuum oven at 120 ° C under a vacuum of 50 mt for 24 hours.

Example 4 [0059] 400 ml of a 0.5 M acetic acid solution (pH = 2.55) was prepared using distilled deionized water (11.6 ml glacial acetic acid was added to 388.4 ml distilled deionized water). A WKI250 water cooling system (Lauda, ​​Westbury, NY, USA) was used to cool the glass reaction vessel to a constant temperature of 4 ° C for one hour. 1.8 g of microfibrillar collagen from bovine tendons (Collagen Matrix Inc, NJ, USA) was added to 320 ml of a 0.5 M acetic acid solution. This suspension was mixed using an IKA Ultra Turrax T18 overhead mixer (IKA Works Inc, Wilmington, NC) at 15,000 rpm for 90 minutes at 4 ° C.

[0060] 40 ml acetic acid solution was mixed with hydroxyapatite (HA) powder (Biotal, UK), exactly 50%, 100% and 200% by weight collagen (0.9, 1.8, and 3.6 g HA) . A 10 mL aliquot of this acetic acid / HA solution was added to the collagen / acetic acid suspension in the cooled reaction vessel after 90 minutes. Then 10 ml aliquots (three in total) were added to the suspension every hour. After the last portion of the acetic acid / HA solution was added, the suspension was stirred for another 60 minutes, resulting in a total mixing time of 330 minutes (five and a half hours). After completion of the mixing step, the suspension was transferred to a clean wide neck beaker and degassed under vacuum at about 4,000 mt for an additional 60 minutes.

[0061] Scaffolding was prepared using a lyophilization (freeze drying) process. An aliquot of 67.5 ml collagen / HA suspension was placed on the wall of the lyophilizer tray provided by the lyophilizer manufacturer (VirTis Co., Gardiner, NY) and made of 304 grade stainless steel. The sample tray had internal dimensions of 127 mm wide x 127 mm long x 38 mm high. The thickness of the plate underlying the tray is 3 mm. The sample tray was placed in the freeze dryer chamber and placed on the cooling shelf of the freeze dryer at 20 ° C.

[0062] The freeze drying process involved cooling the freeze dryer chamber and cooling shelf at a constant cooling rate (0.9 ° C / min) to a final freezing temperature (40 ° C). The shelf and chamber temperature was then kept constant at the final freezing temperature for 60 minutes. The shelf temperature was then raised to 0 ° C in 160 minutes. The ice phase was then sublimated under about 200 mt vacuum at 0 ° C for 17 hours to produce a collagen / HA porous scaffold.

[0063] The porous collagen / HA construct was then placed in a vacuum oven (Fisher IsoTemp 201, Fisher Scientific, Boston, MA) to cross-link collagen in a dehydrothermal crosslinking process. The scaffolds were placed in a vacuum oven at 120 ° C under a vacuum of 50 mt for 24 hours.

[0064] After the DHT crosslinking procedure, the scaffolds were chemically crosslinked using ethyl-3- [3-dimethylaminopropyl] carbodiimide hydrochloride (EDAC) as the crosslinker. EDAC at a concentration of 6 mmol EDAC per gram of scaffolding was mixed in a 5: 2 molar ratio with N-hydroxysuccinimide (EDAC: NHS = 5: 2). The scaffolds were immersed in this EDAC / NHS solution and incubated for 2 hours at room temperature. The scaffolds were then washed twice with phosphate buffered saline (PBS) and incubated in PBS for two hours using an orbital shaker to mix with PBS.

Characteristics of composite scaffolds [0065] For the purposes of this study, all collagen / HA scaffolds produced were compared with a collagen scaffold control scaffold made using the standard protocol used in this research laboratory, in particular in a 0.5 M acetic acid solution and freeze-dried constant cooling rate to a final freezing temperature of 40 ° C.

1. Mechanical stiffness [0066] To ensure survival after implantation into a bone defect, the bone substitute for transplantation must have sufficient internal strength to withstand the forces it is subjected to by bearing the load at the affected site of the defect. The ability to customize osteoconductive bone substitute for transplant with sufficient internal strength to allow implantation into the loaded defect was the main goal of this study. Through the use of composite technology, the extremely biocompatible collagen-based construct is combined with a stronger ceramic hydroxyapatite to develop a bone substitute for a transplant that has the benefits of both and has no disadvantages. All tests were performed using scaffolds hydrated with phosphate buffered saline (PBS). Compression testing of scaffolding samples was carried out using a Zwick strength test machine equipped with a 5-N loading chamber. Samples with a diameter of 8 mm (4 mm high) were cut from the sheets using a leather punch sharpened with a round metal file. The samples were then pre-hydrated with phosphate buffered saline (PBS) for one hour before testing in a 24-well culture plate. The study protocol consisted of two cycles: the preconditioning cycle and the test cycle. For both cycles a pre-load of 0.15 mN was applied and this position was held for one minute. This force was chosen because it was low enough (0.5% load at 10% deformation) to ensure contact with the sample without compressing the sample before testing. The position of the top plate in this preload was used to measure the scaffolding height. Hydrated scaffolds were placed on a dry board, which was then immersed before lowering the upper board. Care was taken to ensure that no bubbles were trapped between the top plate and the scaffolding. For pre-conditioning, the samples were loaded to 5%. For testing, the scaffolds were loaded to 10% and unloaded. A strain rate of 10% per minute was used. After testing, the diameter of the samples was measured at three different locations using a vernier caliper. The module was defined as the slope of the linear fit to the stress-strain curve at 2-5% strain.

[0067] Figure 1 shows the effect of adding HA to the non-crosslinked EDAC scaffold on compression stiffness. It was found that the addition of 50 wt. HA significantly increased the compressive stiffness measured in the strength test with unlimited compression. Compression stiffness increased by almost 300% compared to collagen scaffolds used as control products. Particularly interesting was the effect of acetic acid concentration on the efficiency of HA incorporation into the construct. It is believed to explain the relatively small increase in stiffness when introducing 10 wt. HA in a standard collagen suspension without a corresponding increase in acetic acid concentration. In summary, when adding only 50 wt. HA achieves more than a threefold increase in stiffness, and with a slight adjustment of the acetic acid concentration, maximum benefit can be obtained in terms of structural rigidity by adding relatively small amounts of HA. As a result, this will allow a significant increase in the proportion of HA added by appropriately changing the initial concentration of acetic acid in further studies.

[0068] Figure 2 shows the effect of adding HA to the EDAC cross-linked scaffold on compression stiffness. The starting scaffold, collagen, increases stiffness from 0.2 kPa in Figure 1 to about 1.5 kPa in Figure 2. The addition of HA continues to increase the stiffness of the scaffolds as intended, but with lower amounts of HA, e.g. 50 wt.%, HA , this is obscured by the influence of EDAC crosslinking. However, at 200 wt. HA, a significant increase in stiffness is seen as before. Adding all amounts of HA has been shown to significantly improve the biocompatibility aspects of crosslinked scaffolds, which are discussed in sections 5 and 7 below.

2. Porosity according to the invention [0069] The porosity of the porous scaffold is a measure of the proportion of scaffold volume that is the open pore space, expressed as a percentage. Put simply, this is the percentage of pore volume of the porous construct. The high porosity of the scaffold is required for the diffusion of nutrients / waste materials into / from cells, both in vitro and in vivo. One of the major limitations in developing tissue engineering scaffolds is the issue of core degradation resulting from a lack of nutrient supply and waste removal from the inside of the construct. As a result, constructs often fail after implantation due to necrosis caused by a lack of vascularization in the middle of the scaffolding. One of the great advantages of the collagen-based scaffolds of the present invention is their high porosity. Scaffold porosity was determined by accurately measuring a dry 8 mm scaffold sample 4 mm deep based on the balance sheet ₂ mass. Using the formula for cylinder volume, #h, the density of each sample was calculated by dividing mass by volume. Porosity was calculated according to the formula 100 - [100 (p_ruszt_thatanie / material) l, where p_scaffolding is the density of the sample, ap_{It has}t_eriesj is the weighted density of the ingredients ^{scaffolding (i.e.}scaffolding with 10% of weight by weight^{= [m}collagen ^{+ m}10 wt^{l / [m}collagen^{/ p}collagen ^{+ m}10 wt^{/ p}10 wt^l). [0070] The addition of HA to the constructs reduced the scaffold porosity, but it was negligible in absolute terms, as can be seen in Figure 3.

Thus, the scaffolds of the invention contain a highly porous HA inserted into the construct to improve cell migration to the inside of the scaffold by promoting subsequent cell proliferation. This proved true in the data from in-vivo animal studies presented below.

[0071] In particular, our porosity range within the scaffolds actually produced was from 99.5% for pure collagen to 99% for scaffolds with 200 wt. HA.

3. Mineral structure [0072] The distribution of mineral particles in scaffolds is a difficult parameter / feature for quantitative assessment. This is a feature that is much easier to visualize, and therefore the range of values ​​to be protected is difficult to define. They were visualized using two different ways. The first was the microCT method, which is shown in Figure 6 below. The MicroCT scanner used in this analysis uses x-rays to detect mineralized tissue. As a result, Figure 6 only shows mineral particles in a scaffold with 100 wt. HA. Knowing that collagen is not visible, you can see that the mineral particles are completely and evenly distributed in the scaffolding. Given that the scaffold is 99% hollow, this image provides clear evidence that HA is closely related to collagen fibers. Figure 7 is a 2-dimensional piece of the same scaffold illustrating the arrangement from a different point of view.

[0073] Figures 8 and 9 show the distribution of mineral particles in a scaffold with 50 wt. HA using a separate imaging tool, scanning electron microscopy (SEM). Both images show the same area of ​​interest in scaffolding with 50 wt. HA. Figure 8 shows both phases, collagen and HA. Mineral particles are indistinguishable with the naked eye. However, using X-ray energy dispersion analysis, mineral particles can be detected in an identical area of ​​interest (ROI). This is shown in Figure 9 in the form of white pixels representing mineral particles. In combination with data from microCT, these images clearly prove that the mineral particles are evenly and uniformly distributed in the scaffolding and are closely related to collagen supports.

4. Interconnection between pores [0074] Interconnection between pores is another important feature of the scaffolding, which is very difficult to quantify in scaffoldings consisting primarily of biological material. However, the interconnections between the pores are strongly associated with the permeability of the scaffolding. Permeability is discussed below, but the conductivity of the flow depends on the porosity, pore size and interconnection between the pores. In this regard, permeability indicates interconnection between pores for a given pore size and porosity.

[0075] SEM images of the scaffolds of the invention are provided to illustrate the extremely high levels of interconnection between the pores that can be easily seen. Figure 10 shows a scaffold with 50 wt. HA at 10x magnification and the structure of clearly connected pores on the surface. The same was shown using thin sections taken from such samples. Figure 11 shows the same scaffold with 50 wt. at 100x magnification. This image shows the interconnections between the pores. At this magnification, the pore structure clearly shows interconnection.

5. In-vitro biological activity [0076] The effect of adding HA to the collagen scaffold was assessed by quantifying the osteoblast proliferation MC3T3E1 in scaffolds after 7, 14, 21 and 28 days of incubation. It has been found that the addition of HA to the scaffolds does not have a detrimental effect on cell activity. In fact, the opposite effect was found 28 days after sowing. Increasing the HA share to 200% by weight HA was found to stimulate cell proliferation even more than the control scaffold from pure collagen. Figure 4 shows the absolute numbers of cells retained on the scaffolds.

[0077] As can be seen, scaffolding with 50 wt. HAs maintained significantly fewer cells due to empirical limitations. As a result, Figure 4 does not give the best indication of biological activity. In this regard, Figure 5 illustrates the average number of net cells that remained in the type scaffolds after 28 days. A net decrease of 0.5 million in cell proliferation was observed in pure collagen constructs (about 20% decrease), which is not surprising given that the pure collagen construct is beneficial but not optimized for osteoblast cells. However, the addition of 50% by weight, 100% by weight and 200 wt. HA leads to increased cell proliferation by approximately 500%, 50% and 30% respectively (Figure 4).

6. Permeability of the Invention [0078] The permeability of a porous scaffold generally means the conductivity of the flow under pressure through this porous substrate. The high permeability of the scaffold is essential for the long life of the scaffold in vivo, as it allows cells to migrate to the center of the scaffold and facilitates vascularization in vivo. Scaffolding with 50% by weight, 100% by weight and 200 wt. HA showed significantly increased mean tissue permeability compared to control collagen scaffolds. This was a surprising but positive result of this study because it shows that the addition of HA actually supports the flow of fluid through the scaffold. It is believed that this increase in flow conductivity through a porous scaffold is due to the increased stiffness of the scaffold. The empirical protocol used to quantify scaffold permeability is described in detail in reference [5] (O'Brien et al., 2007, chapter 2.2, pages 7-10).

7. In-vivo animal assay [0079] A small animal assay was conducted to determine the potential of this invention to promote osteogenesis and mineralization of critical size bone loss. Nine Wistar rats were used during the trial. A critical size defect was created in the rat skull vault. One animal was left with an empty defect as a control. The remaining eight animals were divided into groups. Specifically, four cavities were filled with scaffolding with 50 wt. HA, two of which were plated with rat mesenchymal stem cells (MSCs) and two of them were left unsaturated. This served to investigate the potential of the designed tissue in comparison with the available types of scaffolding. The remaining four test cavities were filled with scaffolding with 200 wt. HA and again these four cavities were equally divided into sown and unloaded scaffoldings. After 28 days in the rat skull vault, the animals were sacrificed and the skull vault bones removed. They were treated and analyzed using microCT to investigate the presence of scaffolding in the defect and to observe the effect of scaffolding types on the healing process, osteogenesis and production of the mineralized matrix. The figures in the following chapters show 2-dimensional sections taken from an empty cavity rat. Scraps were frontal sections through the skull bones. The cavities were 5 mm in diameter and were perfectly round. The cross sections shown on the microCT data are schematically shown below in Figure 13.

Empty cavity (Fig. 14) [0080] Data from an empty cavity animal showed that the defect was filled with soft fibrous tissue as part of the healing process. This was expected and was observed in previous animal experiments in our tissue engineering group. In some places in the defect after 28 days of testing, small particles of dense material were observed in the empty defect, but they were rarely observed and not dense enough to indicate a significant healing process in the empty defect sample. Examples are shown on the microCT x-ray images below.

Sown 50 wt. 1 (Fig. 15) [0081] Scaffolds with 50 wt. Cell-seeded HA showed more promising results. Rat mesenchymal stem cells were plated on these scaffolds prior to implantation. As can be seen in the microCT x-ray examples presented below, small heterogeneous pockets of poorly mineralized material were observed not only on the periphery at the junction of the defect with the bone, but also in the center of the scaffolding. The occurrence of these mineralization pockets was much more frequently observed compared to the empty cavity and seemed to be brighter in their intensity compared to the cases observed in the empty cavity.

Sown 50 wt. 2 (Fig. 16) [0082] Our second scaffold with 50 wt. Cell-seeded HA showed significantly better results compared to the first 50% by weight scaffold. HA seeded cells, and thus also with an empty defect. Significant cases of highly mineralized tissue were visible in the area in the defect filled with scaffolding. This was particularly evident around the perimeter where the scaffolding meets the bone. The level of mineralization was not as high as that observed in the surrounding bone, but it was very similar. This is seen as almost identical levels of intensity of the mineralized tissue observed in the defect filled with scaffolding and surrounding bone.

Sown 200 wt. 1 (Fig. 17) [0083] First scaffold with 200 wt. Cell-seeded HA showed significantly better results compared to 50% by weight scaffolds. HA. At almost every point examined in the defect, significant levels of mineralization were observed from the perimeter of the defect to the center of the defect filled with scaffolding. In this sample, the level of mineralization was not as high as in the surrounding bone tissue, as indicated by the relative difference in the intensity of the image of mineralized particles inside the cavity filled with scaffolding.

Sown 200 wt. 2 (Fig. 18) [0084] Second scaffold with 200 wt. Cell-seeded HA showed significantly better results compared to all previous scaffolds. In a significant number of areas studied in the defect, significant levels of mineralization were observed from the perimeter of the defect to the center of the defect filled with scaffolding. Unlike the first scaffolding with 200 wt. HA seeded with cells, the mineralized tissue formed was not of a partial nature, but was continuous throughout the area of ​​the defect filled with scaffolding. Most interestingly, such continuous mineralization was observed in the widest part of the defect. In this sample, the level of mineralization was almost identical to that in the surrounding bone tissue, as indicated by a similar image intensity of the mineralized material inside the cavity filled with scaffolding.

50% wt. I (Fig. 19) [0085] Results from cavities filled with scaffolds without cells for scaffolding with 50 wt. HA were very similar as for scaffolding with 50 wt. HA seeded with cells. Significantly mineralized tissue was observed in defects filled with scaffolding, but it was not continuous. However, the intensity of the mineralized tissue was slightly higher than that observed in cell seeded samples. This was seen in all subsequent unseeded cell samples and indicated that the cell-free construct could function better in vivo.

50% wt. 2 (Fig. 20) [0086] This sample with 50 wt. Cells not shed HA showed similar results to all other samples with 50 wt. HA. After 28 days of trial, evidence of scaffold mineralization initiation was observed both at the perimeter and in the center of the scaffold-filled defect, but the mineralized tissue was not continuous. However, the intensity of the mineralized particles indicated similar mineralization as in the surrounding bone.

200% wt. 1 (Fig. 21) [0087] As observed in samples with 200 wt. HA seeded cells, sample

200 wt. Unsupported HA cells showed a significant level of mineralization around the perimeter and in the center of the defect filled with scaffolding in a significant number of areas in the defect. This mineralized tissue was continuous, in contrast to that observed in samples with 50 wt. HA.

200% wt. 2 (Fig. 22) [0088] As observed in the previous sample with 200 wt. HA not sucked out, this non-sucked sample with 200 wt. HA showed a significant level of mineralization around the perimeter and inside the defect filled with scaffolding in a significant number of areas in the defect. This mineralized tissue was continuous, in contrast to that observed in samples with 50 wt. HA.

[0089] The invention is not limited to the embodiments described herein. As such, the three embodiments described herein represent a small proportion of the total number of scaffold variants that can be made using the same basic manufacturing protocol. It is possible to modify either the ingredients themselves or specific stages of the form or method steps to produce a variable range of constructs optimized for a particular application. Possible modifications include;

[0090] Acetic acid concentration: the acetic acid concentration in the initial collagen suspension can be changed to suit specific applications. Increasing the concentration promotes faster and more homogeneous incorporation of HA particles into the mixed suspension. In addition, this concentration also has a significant impact on both the mechanical properties and the biocompatibility of the scaffolding. These influences are discussed in detail in the chapter "Characteristics of the invention" above. Accordingly, the acetic acid concentration can be varied between 0.05 M and 5M.

[0091] Amount of collagen: the amount of collagen can be changed in the initial collagen suspension. Increasing the amount of collagen leads to increased mechanical rigidity of the obtained scaffolding. It also has a significant impact on the biocompatibility of the scaffolding. Accordingly, the amount of collagen can be changed from 0.5 g / L to 50 g / L acetic acid solution (1/10 and 10 times the standard collagen concentration, respectively).

[0092] Amount of hydroxyapatite: the amount of HA can be varied within a certain range relative to the proportion of collagen in the suspension per scaffold before production. In particular, the amount of HA can be suitably changed from 10 to 1000 weight percent of the amount of collagen used. It was found that increasing the HA content significantly increases the mechanical rigidity of the scaffolding produced.

[0093] Type of hydroxyapatite: The present invention can be made using both sintered, unintended and other forms of HA powder.

[0094] Addition of hydroxyapatite: Both the HA portion volume and injection intervals can be changed to facilitate mixing of the two basic components of the suspension. In general, the intervals between injections can be changed from 30 minutes to 240 minutes. In addition, the portion volume can be varied from 1 ml to 100 ml accordingly. This freedom allows optimization of any particular embodiment of the invention.

[0095] Hydroxyapatite particle size: Typically, HA particle sizes can be varied from 10 nm to 100 μm to suit specific applications.

[0096] Final freezing temperature: the final freezing temperature achieved during the freeze-drying process determines the average pore size in the scaffolds produced. This final freezing temperature can be varied to produce scaffolds with different average pore sizes specific to the particular application or type of cell. Accordingly, the final freezing temperature can be varied from -10 ° C to -70 ° C.

[0097] Freezing surface: The freezing surface between the suspension according to the form and the cooling shelf of the freeze dryer can be changed. The type of freezing surface affects the transfer of energy to / from the suspension / scaffolding and can affect the structure of the pores in the final scaffolding. There are four main options, in particular a vessel with walls of a certain surface, made of metal (1), plastic (2), thin polymer membrane (3) or lack of surface (4).

[0098] Freezing rate: The freezing rate determines the rate of nucleation of ice crystals in the collagen / HA suspension during the freeze drying process and regulates the uniformity of the pore formation process. The cooling rate varies to optimize the freezing process to the different types of available surfaces between the suspension and the freeze dryer cooling shelf (e.g. metal, plastic, none). Usually the freezing rate can be changed from 0.01 ° C / min to 10 ° C / min.

[0099] Annealing: the soaking step can be used in the freeze drying process and allows the formation of pores with an average diameter much larger than the pore size obtainable by changing the final freezing temperature alone. The soak time can be changed from 15 minutes to 36 hours. The longer the soaking time, the greater the final average pore size.

[0100] Scaffold Crosslinking Method: The crosslinking method may be one of several possible techniques, dehydrothermal or chemical. In addition, both techniques can be used sequentially. Specific cross-linking options include glutaraldehyde, carbodiimides (EDAC), microbial transglutaminase (mTgase), dehydrothermal cross-linking (DHT) and ultraviolet (UV) radiation.

[0101] Scaffold Crosslinking Temperature / Concentration: Lyophilized scaffold collagen can be cross-linked by dehydrothermal to increase the mechanical stiffness of the scaffold.

[0102] The crosslinking temperature can be varied from 105 ° C to 180 ° C with a corresponding increase in the stiffness of the form. Additionally, using chemical cross-linking methods, the concentration of the cross-linking solution can be changed to change the degree of chemical cross-linking.

[0103] Scaffold Crosslinking Time: The exposure time to the crosslinker can also be changed to change the degree of crosslinking in the scaffold. It can be changed from 24 to 120 hours to change the final mechanical properties of the scaffolding.

References [0104] [1] Tancred DC, Carr AJ and McCormack BA Development of a new synthetic bone graft. Journal of Materials Science: Materials in Medicine, 9 (12): 819-823, 1998.

[2] Dong JK, Luthy H., Wohlwend A. and Scharer P. Heatpressed ceramics: technology and strength. International Journal of Prosthodontics, 5 (1): 9-16, 1992.

[3] Bailey AJ, Light ND, and Atkins EDT Chemical crosslinking restrictions on models for the molecular organization of the collagen fiber. Nature, 288: 408-410, 1980.

[4] Yannas IV Tissue and Organ Regeneration in Adults. New York: Springer, 2001.

[5] O'Brien FJ, Harley BA, Yannas IV, and Gibson LJ The effect of pore size on cell adhesion in collagen gag scaffolds. Biomaterials, 26: 433-441, 2005.

[6] O'Brien, FJ; Harley, BA; Waller, MA; Yannas, IV; Gibson, LJ and Prendergast, PJ (2007) The effect of pore size on permeability and cell attachment in collagen scaffolds for tissue engineering. Technology and Healthcare Invited Article (15): 3-17.