fabrication of the invention collagen control scaffolds and 10 wt % ha scaffolds were manufactured using the protocol described in embodiment 1 , specifically using an initial acetic acid concentration of 0 . 05m . as the proportion of ha was increased up to 50 wt % ha , homogenous mixing of the two main constituents ( collagen and ha ) became more problematic . an increase in the initial acetic acid concentration was found to alleviate this problem . the effect of this increased acetic acid concentration was investigated using two distinct increases in acetic acid concentration , specifically 0 . 1m and 0 . 5m . these embodiments are described in examples 2 and 3 , respectively . the manufacture of edac crosslinked collagen control scaffolds and composite scaffolds having 50 wt %, 100 wt %, and 200 wt % ha , are described in example 4 . 400 ml of a 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 of distilled , deionized water ). a wk1250 water cooling system ( lauda , westbury , n . y ., usa ) was used to cool a glass reaction vessel to a constant temperature of 4 ° c . for one hour . this reaction vessel was used to blend the scaffold constituents while maintaining the slurry at a constant temperature of 4 ° c . this prevented denaturation of the collagen fibres as a result of heat generation during the blending process . 1 . 8 gm of microfibrillar bovine tendon collagen ( collagen matrix inc ., nj , usa ) was added to 320 ml of the 0 . 05 m acetic acid solution . this suspension was blended using an ika ultra turrax t18 overhead blender ( ika works inc ., wilmington , n . c .) at 15 , 000 rpm for 90 minutes at 4 ° c . 40 ml of the acetic acid solution was mixed with hydroxyapatite ( ha ) powder ( biotal , uk ), specifically 10 % by weight of collagen ( 0 . 18 gm ha ). a 10 ml aliquot of this acetic acid / ha solution was added to the collagen / acetic acid slurry in the cooled reaction vessel at 90 minutes . the method of ha suspension delivery involved a vigourous shaking of the suspension immediately prior to injection ( ensuring a homogenous suspension of the mineral particles ) into the blender vortex centre via syringe . a flexible rubber tube was attached to the syringe nozzle to facilitate injection directly into the blender vortex centre . 10 ml aliquots ( three in total ) were added subsequently to the slurry every hour . after the final aliquot of acetic acid ha solution was added , the slurry was blended for a subsequent 60 minutes , leading to a total blend time of 330 minutes ( five and a half hours ). once the blending stage was completed , the slurry was transferred to a clean , widenecked beaker and vacuum degassed at a pressure of approximately 4000 mtorr for an additional 60 minutes . this stage removed any unwanted air bubbles within the slurry as this would have a detrimental effect on the subsequent freezedrying process . the scaffold was produced using a lyophilisation ( freezedrying ) process . a 67 . 5 ml aliquot of the collagen / ha slurry was placed in a walled freezedryer sample tray supplied by the freezedryer manufacturer ( virtis co ., gardiner , n . y ., usa ) and made from 304 grade stainless steel . the inner sample tray dimensions were 127 mm wide × 127 mm length × 38 mm height . the tray base plate thickness is 3 mm . the sample tray was placed into the freezedryer chamber and placed on the freezedryer cooling shelf at a temperature of 20 ° c . the freezedrying involved the cooling of the freezedryer chamber and cooling shelf at a constant cooling rate ( 0 . 9 ° c ./ min ), based on a previous study , to the final temperature of freezing ( 40 ° c .). the primary determinant of ice crystal morphology during the freezedrying process is the final temperature of freezing . the shelf and chamber temperature was then held constant at the final temperature of freezing for 60 minutes to complete the freezing process . the shelf temperature was then ramped up to 0 ° c . over 160 minutes . the ice phase was then sublimated under a vacuum of approximately 200 mtorr at 0 ° c . for 17 hours to produce the porous collagen / ha scaffold . the porous collagen / ha construct was then placed in a vacuum oven ( fisher isotemp 201 , fisher scientific , boston , mass .) to crosslink the collagen via a dehydrothermal crosslinking process . the scaffolds were placed in the vacuum oven at a temperature of 120 ° c . under a vacuum of 50 mtorr for 24 hours . 400 ml of a 0 . 1 m acetic acid solution ( ph 2 . 9 ) was prepared using distilled , deionised water ( 2 . 32 ml glacial acetic acid was added to 397 . 68 ml of distilled , deionized water ). a wk1250 water cooling system ( lauda , westbury , n . y ., usa ) was used to cool a glass reaction vessel to a constant temperature of 4 ° c . for one hour . 1 . 8 gm of microfibrillar bovine tendon collagen ( collagen matrix inc ., nj , usa ) was added to 320 ml of the 0 . 1 m acetic acid solution . this suspension was blended using an ika ultra turrax t18 overhead blender ( ika works inc ., wilmington , n . c .) at 15 , 000 rpm for 90 minutes at 4 ° c . 40 ml of the acetic acid solution was mixed with hydroxyapatite ( ha ) powder ( biotal , uk ), specifically 50 % by weight of collagen ( 0 . 9 gm ha ). a 10 ml aliquot of this acetic acid / ha solution was added to the collagen / acetic acid slurry in the cooled reaction vessel at 90 minutes . 10 ml aliquots ( three in total ) were added subsequently to the slurry every hour . after the final aliquot of acetic acid ha solution was added , the slurry was blended for a subsequent 60 minutes , leading to a total blend time of 330 minutes ( five and a half hours ). once the blending stage was completed , the slurry was transferred to a clean , widenecked beaker and vacuum degassed at a pressure of approximately 4000 mtorr for an additional 60 minutes . the scaffold was produced using a lyophilisation ( freezedrying ) process . a 67 . 5 ml aliquot of the collagen / ha slurry was placed in a walled freezedryer sample tray supplied by the freezedryer manufacturer ( virtis co ., gardiner , n . y ., usa ) and made from 304 grade stainless steel . the inner sample tray dimensions were 127 mm wide × 127 mm length × 38 mm height . the tray base plate thickness is 3 mm . the sample tray was placed into the freezedryer chamber and placed on the freezedryer cooling shelf at a temperature of 20 ° c . the freezedrying process involved the cooling of the freezedryer chamber and cooling shelf at a constant cooling rate ( 0 . 9 ° c ./ min ), to the final temperature of freezing ( 40 ° c .). the shelf and chamber temperature was then held constant at the final temperature of freezing for 60 minutes . the shelf temperature was then ramped up to 0 ° c . over 160 minutes . the ice phase was then sublimated under a vacuum of approximately 200 mtorr at 0 ° c . for 17 hours . the porous collagen / ha construct was then placed in a vacuum oven ( fisher isotemp 201 , fisher scientific , boston , mass .) to crosslink the collagen via a dehydrothermal crosslinking process . the scaffolds were placed in the vacuum oven at a temperature of 120 ° c . under a vacuum of 50 mtorr for 24 hours . 400 ml of a 0 . 5 m acetic acid solution ( ph 2 . 55 ) was prepared using distilled , deionised water ( 11 . 6 ml glacial acetic acid was added to 388 . 4 ml of distilled , deionized water ). a wk1250 water cooling system ( lauda , westbury , n . y ., usa ) was used to cool a glass reaction vessel to a constant temperature of 4 ° c . for one hour . 1 . 8 gm of microfibrillar bovine tendon collagen ( collagen matrix inc ., nj , usa ) was added to 320 ml of the 0 . 5 m acetic acid solution . this suspension was blended using an ika ultra turrax t18 overhead blender ( ika works inc ., wilmington , n . c .) at 15 , 000 rpm for 90 minutes at 4 ° c . 40 ml of the acetic acid solution was mixed with hydroxyapatite ( ha ) powder ( biotal , uk ), specifically 50 %, 100 %, and 200 %, by weight of collagen ( 0 . 9 , 1 . 8 , and 3 . 6 gm ha ). a 10 ml aliquot of this acetic acid / ha solution was added to the collagen / acetic acid slurry in the cooled reaction vessel at 90 minutes . 10 ml aliquots ( three in total ) were subsequently added to the slurry every hour . after the final aliquot of acetic acid ha solution was added , the slurry was blended for a subsequent 60 minutes , leading to a total blend time of 330 minutes ( five and a half hours ). once the blending stage was completed , the slurry was transferred to a clean , widenecked beaker and vacuum degassed at a pressure of approximately 4000 mtorr for an additional 60 minutes . the scaffold was produced using a lyophilisation ( freezedrying ) process . a 67 . 5 ml aliquot of the collagen / ha slurry was placed in a walled freezedryer sample tray supplied by the freezedryer manufacturer ( virtis co ., gardiner , n . y ., usa ) and made from 304 grade stainless steel . the inner sample tray dimensions were 127 mm wide × 127 mm length × 38 mm height . the tray base plate thickness is 3 mm . the sample tray was placed into a freezedryer chamber and placed on the freezedryer cooling shelf at a temperature of 20 ° c . the freezedrying process involved the cooling of the freezedryer chamber and cooling shelf at a constant cooling rate ( 0 . 9 ° c ./ min ), to the final temperature of freezing ( 40 ° c .). the shelf and chamber temperature was then held constant at the final temperature of freezing for 60 minutes . the shelf temperature was then ramped up to 0 ° c . over 160 minutes . the ice phase was then sublimated under a vacuum of approximately 200 mtorr at 0 ° c . for 17 hours to produce the porous collagen / ha scaffold . the porous collagen / ha construct was then placed in a vacuum oven ( fisher isotemp 201 , fisher scientific , boston , mass .) to crosslink the collagen via a dehydrothermal crosslinking process . the scaffolds were placed in the vacuum oven at a temperature of 120 ° c . under a vacuum of 50 mtorr for 24 hours . 400 ml of a 0 . 5 m acetic acid solution ( ph 2 . 55 ) was prepared using distilled , deionised water ( 11 . 6 ml glacial acetic acid was added to 388 . 4 ml of distilled , deionized water ). a wk1250 water cooling system ( lauda , westbury , n . y ., usa ) was used to cool a glass reaction vessel to a constant temperature of 4 ° c . for one hour . 1 . 8 gm of microfibrillar bovine tendon collagen ( collagen matrix inc ., nj , usa ) was added to 320 ml of the 0 . 5 m acetic acid solution . this suspension was blended using an ika ultra turrax t18 overhead blender ( ika works inc ., wilmington , n . c .) at 15 , 000 rpm for 90 minutes at 4 ° c . 40 ml of the acetic acid solution was mixed with hydroxyapatite ( ha ) powder ( biotal , uk ), specifically 50 %, 100 %, and 200 %, by weight of collagen ( 0 . 9 , 1 . 8 , and 3 . 6 gm ha ). a 10 ml aliquot of this acetic acid / ha solution was added to the collagen / acetic acid slurry in the cooled reaction vessel at 90 minutes . 10 ml aliquots ( three in total ) were subsequently added to the slurry every hour . after the final aliquot of acetic acid ha solution was added , the slurry was blended for a subsequent 60 minutes , leading to a total blend time of 330 minutes ( five and a half hours ). once the blending stage was completed , the slurry was transferred to a clean , widenecked beaker and vacuum degassed at a pressure of approximately 4000 mtorr for an additional 60 minutes . the scaffold was produced using a lyophilisation ( freezedrying ) process . a 67 . 5 ml aliquot of the collagen / ha slurry was placed in a walled freezedryer sample tray supplied by the freezedryer manufacturer ( virtis co ., gardiner , n . y ., usa ) and made from 304 grade stainless steel . the inner sample tray dimensions were 127 mm wide × 127 mm length × 38 mm height . the tray base plate thickness is 3 mm . the sample tray was placed into a freezedryer chamber and placed on the freezedryer cooling shelf at a temperature of 20 ° c . the freezedrying process involved the cooling of the freezedryer chamber and cooling shelf at a constant cooling rate ( 0 . 9 ° c ./ min ), to the final temperature of freezing ( 40 ° c .). the shelf and chamber temperature was then held constant at the final temperature of freezing for 60 minutes . the shelf temperature was then ramped up to 0 ° c . over 160 minutes . the ice phase was then sublimated under a vacuum of approximately 200 mtorr at 0 ° c . for 17 hours to produce the porous collagen / ha scaffold . the porous collagen / ha construct was then placed in a vacuum oven ( fisher isotemp 201 , fisher scientific , boston , mass .) to crosslink the collagen via a dehydrothermal crosslinking process . the scaffolds were placed in the vacuum oven at a temperature of 120 ° c . under a vacuum of 50 mtorr for 24 hours . following the dht crosslinking procedure , the scaffolds were chemically crosslinked using ethyl - 3 -[ 3 - dimethylaminopropyl ] carbodiimide hydrochloride ( edac ) as the crosslinking agent . edac at a concentration of 6 mmol edac per gram of scaffold was mixed in a 5 : 2 molar ratio with n - hydroxysuccinimide ( edac : nhs = 5 : 2 ). scaffolds were immersed in this edac / nhs solution and incubated for 2 hours at room temperature . subsequently scaffolds were rinsed twice using phosphate buffered saline ( pbs ) and incubated in pbs for two hours using an orbital shaker to agitate the pbs . for the purpose of this study , all fabricated collagen / ha scaffolds were compared against a control scaffold made of collagen , fabricated using the standard protocol used within this research laboratory , specifically in 0 . 5m acetic acid solution and lyophilised at a constant cooling rate to a final freezing temperature of 40 ° c . to ensure survival once implanted into a bone defect , a bone graft substitute must possess sufficient intrinsic strength to withstand the forces it is subjected to through load bearing within the affected defect site . the ability to custom fabricate an osteoconductive bone graft substitute with sufficient intrinsic strength to allow implantation into a load - bearing defect was the primary goal of this study . through the use of composite technology , an extremely biocompatable collagen - based construct is combined with the stronger ceramic hydroxyapatite to develop a bone graft substitute with all of the advantages of both materials with none of their disadvantages . all tests were conducted on scaffolds hydrated with phosphate buffered saline ( pbs ) solution . compression testing of scaffold samples was carried out using a zwick mechanical testing machine fitted with a 5 - n load cell . samples of 8 mm diameter ( 4 mm high ) were cut from the sheets using a leather punch sharpened by a round metal file . the samples were then pre - hydrated with phosphate buffered saline ( pbs ) one hour prior to testing in a 24 - well culture plate . the testing protocol consisted of two cycles : a precondition cycle and a test cycle . for both cycles a preload of 0 . 15 mn was applied and the position was held for one minute . this force was selected as it was low enough ( 0 . 5 % of the load at 10 % strain ) to ensure contact with the sample without compressing the sample before the test . the position of the upper platen at this preload was used to measure the height of the scaffold . hydrated scaffolds were placed on a dry platen which was then submerged prior to lowering the upper platen . care was taken to insure that no bubbles were trapped between the upper platen and the scaffold . for preconditioning , 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 separate locations using a vernier calipers . the modulus was defined as the slope of a linear fit to the stress - strain curve over 2 - 5 % strain fig1 shows the effect of adding ha to the non - edac crosslinked scaffold on compressive stiffness . the addition of 50 wt % ha was found to significantly increase the compressive stiffness measured by mechanical testing in unconfined compression . an increase in compressive stiffness of nearly 300 % was seen relative to collagen scaffold product controls . of particular interest was the effect of acetic acid concentration on the efficiency of ha incorporation within the construct . this is believed to explain the relatively small increase in stiffness by incorporating 10 wt % ha in the standard collagen slurry without increasing the acetic acid concentration accordingly . in conclusion , with as little as 50 wt % ha added , a greater than three fold increase in stiffness is achieved and by small adjustments to the acetic acid concentration , maximum gains in construct stiffness can be achieved with the addition of relatively small amounts of ha . consequently , this will allow significant increases in the proportion of ha added by altering the initial acetic acid concentration accordingly in future studies . fig2 shows the effect of adding ha to the edac crosslinked scaffold on compressive stiffness . the baseline scaffold , collagen , increases in stiffness from 0 . 2 kpa in fig1 to about 1 . 5 kpa in fig2 . addition of ha still increases the stiffness of the scaffolds as intended but at lower amounts of ha , such as 50 wt % ha , the effect of the edac crosslinking overshadows this . however , at 200 wt % ha , a significant increase in stiffness is seen as before . the addition of all amounts of ha was shown to significantly improve biocompatibility aspects of the crosslinked scaffolds and these are discussed in sections 5 and 7 below . the porosity of a porous scaffold is a measure of the proportion of the scaffold volume composed of open , porous space expressed as a percentage . in simpler terms , it is the percentage pore volume of a porous construct . a high scaffold porosity is required for diffusion of nutrients / waste materials to / from cells both in vitro and in vivo . one of the main constraints in the development of engineered tissue scaffolds has been the issue of core degradation , arising from lack of nutrient delivery and waste removal from the centre of the construct . as a result , constructs often fail once implanted due to avascular necrosis in the centre of the scaffold . one of the big advantages of the collagen - based scaffolds of the present invention is their high porosity . scaffold porosity was determined by the precise measurement of a dry 8 mm , 4 mm deep scaffold sample using a mass balance . using the formula for the volume of a cylinder , πr 2 h , the density of each sample was calculated by dividing the mass by the volume . porosity was calculated using the formula 100 −[ 100 ( ρ scaffold / ρ material )] where ρ scaffold is the density of a given sample and ρ material is the weighted density of the scaffold constituents ( i . e . ρ 10wt % hascaffold =[ m collagen + m 10wt % ha ]/[ m collagen / ρ collagen + m 10wt % ha / ρ 10wt % ha ]) the addition of ha to the constructs resulted in a decrease in scaffold porosity but this was negligible in absolute terms as can be seen in fig3 . therefore , the scaffolds of the invention incorporate ha into the construct while retaining a very high porosity to improve cellular migration into the centre of the scaffold to encourage subsequent cell proliferation . this was shown to be true in the in - vivo animal study data shown below . specifically , our range of porosities within scaffolds actually produced ranges from 99 . 5 % for pure collagen down to 99 % for the 200 wt % ha scaffolds . the distribution of mineral particles throughout the scaffolds is a difficult parameter / attribute to quantify . it is an attribute that is much easier to visualize and therefore , a range of values to protect is difficult to define . it was visualized using two different methods . the first was microct which is shown in fig6 below . the microct scanner used in this analysis uses x - rays to detect mineralised tissue . consequently , fig6 shows only the mineral particles within a 100 wt % ha scaffold . knowing that the collagen is not visible , one can see that the mineral particles are completely and evenly distributed throughout the scaffold . given that the scaffold is 99 % empty , this image shows conclusive evidence that the ha is intimately associated with the collagen fibres . fig7 shows a 2 - dimensional slice of the same scaffold illustrating the distribution from another viewpoint . fig8 and 9 illustrate the distribution of mineral particles throughout a 50 wt % ha scaffold using a distinct imaging tool , scanning electron microscopy ( sem ). both images show the same region of interest within a 50 wt % ha scaffold . fig8 shows both phases of collagen and ha . the mineral particles are indistinguishable to the naked eye . however , using energy dispersive x - ray analysis , the mineral particle can be detected for the identical region of interest ( roi ). this is shown in fig9 as white pixels representing mineral particles . in conjunction with the microct data , these images prove conclusively that the mineral particles are evenly and homogenously distributed throughout the scaffold and are intimately associated with the collagen struts . pore interconnectivity is another important scaffold attribute that is very difficult to define quantitatively in scaffolds composed primarily of biological material . however , pore interconnectivity is strongly related to scaffold permeability . permeability is discussed below but the conductivity of flow is dependant on porosity , pore size and pore interconnectivity together . consequently , permeability gives an indication of pore interconnectivity when pore size and porosity are defined . sem images of scaffolds of the invention are provided to illustrate the extremely high levels of pore interconnectivity which can be easily seen . fig1 shows a 50 wt % ha scaffold at 10 times magnification and the obviously interconnected pore structure at the surface . this has been shown to be identical using thin sections taken from such samples . fig1 shows the same 50 wt % scaffold at 100 times magnification . this image illustrates the pore interconnectivity . at this magnification , the pore structure is conclusively interconnected . the effect of the addition of ha to the collagen scaffold was assessed by quantifying the proliferation of mc3t3e1 osteoblasts in the scaffolds after 7 , 14 , 21 and 28 days incubation . the addition of ha to the scaffolds was found to have no detrimental effect on cellular activity . in fact , the opposite was found to be true at 28 days post seeding . increasing the proportion of ha to 200 wt % ha was found to encourage cell proliferation even further than the control pure collagen scaffold . fig4 shows the absolute cell numbers retained on the scaffolds . as can be seen the 50 wt % ha scaffolds were retained a significantly lower number of cells due to empirical restrictions . consequently , fig4 does not give the best indication of bioactivity . consequently , fig5 illustrates the net mean cell number left within the scaffold types at 28 days . a net decrease of 0 . 5 million in cell proliferation is seen in the pure collagen constructs ( approximately 20 % decrease ), which is not unsurprising given that the pure collagen construct is favourable but not optimized for osteoblast cells . however , the addition of 50 wt %, 100 wt % and 200 wt % ha results in a increased cellular proliferation of approximately 500 %, 50 % and 30 % respectively ( fig4 ). the permeability of a porous scaffold is essentially the flow conductivity under pressure through that porous medium . a high scaffold permeability is essential for long term viability of the scaffold in vivo as it allows cells to migrate to the scaffold centre and facilitates vascularisation in vivo . the 50 wt %, 100 wt % and 200 ha scaffolds exhibited a significantly increased mean tissue permeability relative to the control collagen scaffolds . this was a surprising but positive finding of this study as it shows that the addition of ha actually aids the flow of fluid throughout the scaffold . it is believed that this increase in flow conductivity through the porous scaffold is due to the increased scaffold rigidity . the empirical protocol used to quantify scaffold permeability is described in detail in reference [ 5 ] ( o &# 39 ; brien et al , 2007 , section 2 . 2 , pages 7 - 10 ). a small animal trial was carried out to determine the potential of this invention to encourage osteogenesis and mineralisation of a critically sized defect in bone . nine wistar rats were used during the trial . a critically sized defect was created in the rat calvaria . one animal was left with an empty defect as a control . the other eight animals were split into groups . specifically , four of the defects were filled with 50 wt % ha scaffolds , two of which were seeded with rat mesenchymal stem cells ( msc ) cells and two of which that were left unseeded . this was to investigate the potential of tissue engineered versus off the shelf scaffold types . the remaining four sample defects were filled with 200 wt % ha scaffolds and once again , these four were split evenly between seeded and unseeded scaffolds . after 28 days within the rat calvaria , the animals were sacrificed and the calvarial bones were removed . these were process and analysed using microct to investigate the presence of scaffold within the defect and to observe the effect of scaffold types on the healing process , osteogenesis and mineralised matrix production . the figures in the following sections show 2 - dimensional slices taken from the empty defect rat . slices were coronal sections through the calvarial bone . the defects were 5 mm in diameter and perfectly circular . the sections shown in the microct data are represented in the schematic below in fig1 . the empty defect animal data showed the defect was filled in with soft fibrous tissue as part of the healing process . this was expected and has been seen in previously conducted animal trials within our tissue engineering group . at some points throughout the defect after the 28 day trial , small particles of dense material were observed within the empty defect but these were seldom observed and not sufficiently dense to indicate significant healing within the empty defect sample . examples of these are shown in the microct x - ray images shown below . the 50 wt % ha cell - seeded scaffolds showed more promising results . these scaffolds were seeded with rat mesenchymal stem cells prior to implantation . as can be seen from the representative microct x - ray images below , small heterogeneous pockets of mildly mineralized material were seen not only at the periphery of the defect - bone interface but also in the centre of the scaffolds . there were a significantly higher number of instances of these pockets of mineralization seen relative to the empty defect data and they appeared to be brighter in their intensity relative to the instances seen in the empty defect . our second cell seeded 50 wt % ha scaffold showed much improved results relative to the first cell seeded 50 wt % ha scaffold and consequently the empty defect . significant instances of heavily mineralised tissue were seen at points throughout the scaffold - filled defect . this was especially evident at the periphery of the scaffold - bone interface . the level of mineralisation was not high as that seen in the surrounding bone but was very similar . this can be seen as the nearly identical intensity levels of the mineralised tissue seen within the scaffold - filled defect and that of the surrounding bone . the first cell - seeded 200 wt % ha scaffold showed significantly improved results compared with the 50 wt % ha scaffolds . at nearly every point examined throughout the defect , significant levels of mineralisation were seen from the periphery of the defect all the way into the centre of the scaffold - filled defect . in this sample , the mineralisation level was not as high as that of the surrounding bone tissue indicated by the relative difference in the image intensity of the mineralised particles within the scaffold - filled defect . the second cell - seeded 200 wt % ha scaffold showed significantly improved results compared with all previous scaffolds . at a significant number of areas examined throughout the defect , significant levels of mineralisation were seen from the periphery of the defect all the way into the centre of the scaffold - filled defect . unlike the first 200 wt % ha cell - seeded scaffolds , the mineralised tissue formed was not particulate in nature but was continuous across the scaffold - filled defect . most interestingly , this continuous mineralisation was seen at the widest part of the defect . in this sample , the mineralisation level was nearly identical to that of the surrounding bone tissue indicated by the similar image intensity of the mineralised material within the scaffold - filled defect . the cell - free scaffold - filled defect results for the 50 wt % ha scaffolds were very similar to the cell - seeded 50 wt % ha scaffolds . significantly mineralised tissue was seen throughout the scaffold - filled defects but was not continuous in nature . however , the intensity of the mineralised tissue was marginally greater than that seen in the cell - seeded samples . this was seen in all cell unseeded samples that follow and indicated that a cell - free construct may perform better in vivo . this 50 wt % ha cell - unseeded sample showed similar results to all other 50 wt % ha samples . after the 28 day trial , evidence of the beginning of scaffold mineralisation was seen at both the periphery and the centre of the scaffold - filled defect but the mineralised tissue was not continuous . however , the intensity of the mineralised particles indicated a mineralisation similar to that of the surrounding bone . as seen in the 200 wt % ha cell - seeded samples , the unseeded 200 wt % ha sample showed significant levels of mineralisation at the periphery and the centre of the scaffold - filled defect at a significant number of areas throughout the defect . this mineralised tissue was continuous in nature unlike that seen in the 50 wt % ha samples . as seen in the previous 200 wt % ha cell - unseeded sample , this unseeded 200 wt % ha sample showed significant levels of mineralisation at the periphery and the centre of the scaffold - filled defect at a significant number of areas throughout the defect . this mineralised tissue was continuous in nature unlike that seen in the 50 wt % ha samples . the invention is not limited to the embodiments herein described . as such , the three embodiments of the invention described herein represent a small proportion of the total number of scaffold variants possible using the same core fabrication protocol . either the constituents themselves or specific embodiment stages or process steps can be varied to produce a varied range of constructs , optimised for application specific use . the variations possible include ; acetic acid concentration : the concentration of the acetic acid within the initial collagen slurry can be altered to suit specific applications . increasing the concentration encourages a more rapid and homogenous integration of the ha particles within the blending slurry . additionally , this concentration also has a significant effect on both the mechanical properties and the biocompatability of the scaffold . these effects are discussed in detail in the “ invention characterisation ” section above . suitably , the acetic acid concentration can be varied between 0 . 05m and 5m . collagen quantity : the collagen quantity can be varied within the initial collagen slurry . increasing the collagen quantity results in an increased mechanical stiffness of the resulting scaffold . this also has a significant effect on scaffold biocompatability . suitably , the collagen quantity can vary from 0 . 5 g / l up to 50 g / l of acetic acid solution ( 1 / 10 and 10 times standard collagen concentration respectively ). hydroxyapatite quantity : the quantity of ha can be varied within a specified range relative to the proportion of collagen within the scaffold slurry prior to manufacture . specifically , the quantity of ha can suitably vary from 10 to 1000 weight percent of the quantity of collagen used . increasing the ha content was found to significantly increase the mechanical stiffness of the manufactured scaffold . hydroxyapatite type : the present invention can be manufactured using both sintered , unsintered , and other forms of the ha powder . hydroxyapatite addition : both the ha aliquot volume and injection interval can be varied to facilitate mixing of the two primary constituent of the slurry . generally , the injection interval can be varied from 30 minutes up to 240 minutes . additionally , the aliquot volume can suitably be varied from 1 ml up to 100 ml . such freedom facilitates optimisation of any particular invention embodiment . hydroxyapatite particle size : typically , the ha particle size can be varied from 10 nm up to 100 μm to suit specific applications . final freezing temperature : the final freezing temperature reached during the freezedrying process determines the mean pore size within the manufactured scaffolds . this final freezing temperature can be varied to produce scaffolds with various mean pore sizes specific to a specific application or cell type . suitably , the final freezing temperature can be varied from − 10 ° c . down to − 70 ° c . freezing interface : the freezing interface placed between the embodiment slurry and the freezedryer cooling shelf can be varied . the type of freezing interface affects heat energy transmission to / from the slurry / scaffold and can alter the pore structure of the final scaffold . four main options are available , specifically a walled vessel of defined interface made of either metal ( 1 ), plastic ( 2 ), a thin polymeric membrane ( 3 ) or no interface ( 4 ). freezing rate : the freezing rate determines the rate of ice crystal nucleation within the collagen / ha slurry during the freezedrying process and controls the homogeneity of the pore formation process . the cooling rate is varied to optimise the freezing process for the various types of interfaces available between the slurry and the freezedryer cooling shelf ( e . g . metal , plastic , none ). typically , the freezing rate can be varied from 0 . 01 ° c ./ min up to 10 ° c ./ min . annealing : an annealing stage an be employed during the freezedrying process and allows the creation of pores with an average diameter significantly greater than pore sizes attainable by varying the final freezing temperature alone . the annealing time can be varied from 15 minutes to 36 hours . the longer the annealing time , the larger the final average pore size . scaffold crosslinking method : the crosslinking method can be one of a possible number of techniques , either dehydrothermal or chemical in nature . additionally , both techniques can be employed consecutively . specific crosslinking options include glutaraldehyde , carbodiimides ( edac ), microbial transglutaminase ( mtgase ), dehydrothermal crosslinking ( dht ) and ultraviolet radiation ( uv ). scaffold crosslinking temperature / concentration : the freezedried scaffold collagen can be crosslinked via dehydrothermal means to increase the scaffold mechanical stiffness . the crosslinking temperature can be varied from 105 ° c . up to 180 ° c . with a corresponding increase in embodiment stiffness . additionally , when using chemical crosslinking methods , the concentration of the crosslinking solution can be varied to alter the extent of chemical crosslinking . scaffold crosslinking duration : the crosslink exposure time can also be varied to alter the extent of the crosslinking process throughout the scaffold . this can be varied between 24 and 120 hours to alter the final mechanical specification of the scaffold . tancred d . c ., carr a . j ., and mccormack b . a . development of a new synthetic bone graft . journal of materials science : materials in medicine , 9 ( 12 ): 819 - 823 , 1998 . dong j . k ., luthy h ., wohlwend a ., and scharer p . heatpressed ceramics : technology and strength . international journal of prosthodontics , 5 ( 1 ): 9 - 16 , 1992 . bailey a . j ., light n . d ., and atkins e . d . t . chemical crosslinking restrictions on models for the molecular organisation of the collagen fibre . nature , 288 : 408 - 410 , 1980 . yannas i . v . tissue and organ regeneration in adults . new york : springer , 2001 . o &# 39 ; brien f . j ., harley b . a ., yannas i . v ., and gibson l . j . the effect of pore size on cell adhesion in collagen gag scaffolds . biomaterials , 26 : 433 - 441 , 2005 . o &# 39 ; brien , f . j . ; harley , b . a . ; waller , m . a . ; yannas , i . v ; gibson , l . j and prendergast , p . j . ( 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 .
a process for producing a collagen / hydroxyapatite composite scaffold comprises the steps of forming a homogenous suspension of collagen and ha in an acidic solution , lyophilising the suspension until a desired final freezing temperature is reached to produce the composite scaffold , and optionally cross - linking the composite scaffold , wherein the ratio of ha to collagen is at least 1 : 10 . also provided is a collagen / hydroxyapatite composite scaffold comprising a homogenous distribution of hydroxyapatite within a porous , crosslinked , collagen matrix , wherein the ratio of ha to collagen is at least 1 : 10 . suitably , the composite scaffold has a porosity of at least 99 % , and a compressive stiffness of at least 0 . 3 kpa . composite scaffolds of the invention may be used to provide osteoconductive bone implants and tissue engineering implants .