**2. Cellulose for medical applications and as a tissue engineering matrix**

Cellulose - A Biomaterial with Cell-Guiding Property 85

Cellulose is non-toxic and has good biocompability, therefore, it offers several possibilities in medical applications. Cellulose and its derivatives are used, among other things, as coating materials for drugs, additives of pharmaceutical products, blood coagulant, supports for immobilized enzymes, artificial kidney membranes, stationary phases for optical resolution, in wound care and as implant material and scaffolds in tissue engineering

Cellulose sponges can be manufactured by adding supportive strengthening fibres (8-10 mm long cotton fibres; about 20% of the weight of the cellulose) and sodium sulphate crystals as pore forming material to a cellulose viscose (sodium xantogenate) solution (4-6 g cellulose/100 g viscose). The cellulose is regenerated by heating the solution in a water bath after which the sponge is washed with hot water, treated with a dilute acid and sodium hypochlorite bleaching solution, and finally washed repetitively in distilled water before drying and sterilisation 20, 21. When inserted subcutaneously, a vital and well vascularised repair tissue, called granulation tissue, grows rapidly into this cellulose sponge. Due to this good granulation tissue formation ability, cellulose sponges have been used in experimental surgery for decades 6, 7, 22 and the subcutaneous implantation of the cellulose sponge is widely accepted method for wound healing (see box 2) studies 8, 23. Several cellulose products for wound healing purposes (e.g. Cellospon®, Cellstick®, Sponcal®, Visella®, and Absorpal®) are commercially available. These products are made from the sponge form viscose cellulose and have homogenous porous structure, characterized by thin pore walls with one or more inter-pore openings. They are elastic and can be compressed and expanded repeatedly with no damage to their internal structure, hence providing free entrance for the invading cells to the inner parts of the sponge 24.

Host reactions following implantation of biomaterials include injury, blood-material interactions, provisional matrix formation, inflammation, granulation tissue development, foreign body reaction, and fibrosis/fibrous capsule development 25. When implanted subcutaneously, a blood-material interaction occurs with protein adsorption to the cellulose sponge and a blood-based transient provisional matrix, a blood clot; is formed on and around the sponge. The platelets, originated from the injured blood vessels, not only participate to haemostasis but also liberate bioactive agents like cytokines and growth factors that will attract inflammatory and phagocytosing cells. The first cells to arrive are polymorphonuclear leucocytes, i.e. neutrophils, which are characteristic for the acute inflammatory response. These cells secrete pro-inflammatory cytokines that, in turn, attract circulating monocytes, which are activated and converted in the tissue to macrophages that kill bacterial pathogens, scavenge tissue debris and destroy remaining neutrophils. Biomaterial surface adherent macrophages can also fuse to form multinucleated foreign body giant cells. In their attempt to phagocytose the biomaterial, adherent macrophages become active 25. By releasing a variety of chemotactic, neovasculogenic and growth factors that stimulate cell migration, proliferation and formation of new blood vessels and tissue matrix, macrophages mediate the transition from the inflammatory phase to the

3, 12.

**2.1. Regenerated cellulose** 

Cellulose, the most common organic compound on Earth, is degraded by microbial enzymes. Animal cells cannot cleave the β(1→4)-bond between the two glucose moieties in cellulose. Thus, cellulose degradation in tissues takes place by a slow non-enzymatic hydrolysis of the β(1→4)-bond and therefore cellulose can be regarded as an almost stable matrix. Despite this, cellulose and its derivatives are well tolerated by cells and tissues and induce a moderately strong foreign body reaction in the tissue [3-8].

### **BOX 1. ADULT BONE MARROW-DERIVED STEM CELLS**

Adult stem cells are immature cells, dispersed in tissues throughout the body after development. Like all stem cells, they are capable of either making identical copies of themselves or to differentiate depending on their local environment into mature cell types with characteristic morphology and function. Stem cells usually generate an intermediate, partly differentiated, cell type, called precursor or progenitor cell, before they achieve their fully differentiated state. Adult stem cells are rare, however. Their primary functions are to replenish dying cells, and with limitations, to regenerate damaged tissues.

The best characterised adult stem cells are those found in the bone marrow, which provides a unique niche for haematopoietic stem cells (HSCs) and the mesenchymal stem or stromal cells (MSCs). HSCs are responsible for the production and replacement of all blood cells during the entire lifetime 13. The earliest haematopoietic precursor, the haemangioblast, is not only a precursor of haematopoietic cell lineages but also of cells that line all blood vessels and lymphatics, namely the endothelial cells 14.

Mesenchymal stromal cells are a heterogenous population of stem/progenitor cells able to differentiate into several cell types such as chondrocytes, osteocytes, fibroblasts, myocytes, adipocytes, epithelial and neuron-like cells. When stimulated by specific signals, these cells can be released from their niche in the bone marrow into circulation and recruited to the target tissues where they undergo *in situ* differentiation and contribute to tissue homeostasis and repair 15. MSCs also secrete factors that promote survival and differentiation of endogenous cells as well as angiogenetic factors essential for blood vessel formation. MSCs possess remarkable immunosuppressive properties and can inhibit the proliferation and function of the major immune cell population 16, 17 as well as antimicrobial properties 18. Furthermore, these multipotential stromal stem and progenitor cells at different stages of maturation contribute to the formation of HSC stem cell niche and play a critical role in haematopoiesis 19.The characteristic of MSCs makes these cells exceptionally suitable for various therapeutic possibilities such as supporting tissue regeneration, correcting inherited disorders, dampening chronic inflammation, and delivering biological agents 15.

Cellulose is non-toxic and has good biocompability, therefore, it offers several possibilities in medical applications. Cellulose and its derivatives are used, among other things, as coating materials for drugs, additives of pharmaceutical products, blood coagulant, supports for immobilized enzymes, artificial kidney membranes, stationary phases for optical resolution, in wound care and as implant material and scaffolds in tissue engineering 3, 12.

## **2.1. Regenerated cellulose**

84 Cellulose – Medical, Pharmaceutical and Electronic Applications

induce a moderately strong foreign body reaction in the tissue [3-8].

replenish dying cells, and with limitations, to regenerate damaged tissues.

line all blood vessels and lymphatics, namely the endothelial cells 14.

delivering biological agents 15.

**BOX 1. ADULT BONE MARROW-DERIVED STEM CELLS** 

**2. Cellulose for medical applications and as a tissue engineering matrix** 

Cellulose, the most common organic compound on Earth, is degraded by microbial enzymes. Animal cells cannot cleave the β(1→4)-bond between the two glucose moieties in cellulose. Thus, cellulose degradation in tissues takes place by a slow non-enzymatic hydrolysis of the β(1→4)-bond and therefore cellulose can be regarded as an almost stable matrix. Despite this, cellulose and its derivatives are well tolerated by cells and tissues and

Adult stem cells are immature cells, dispersed in tissues throughout the body after development. Like all stem cells, they are capable of either making identical copies of themselves or to differentiate depending on their local environment into mature cell types with characteristic morphology and function. Stem cells usually generate an intermediate, partly differentiated, cell type, called precursor or progenitor cell, before they achieve their fully differentiated state. Adult stem cells are rare, however. Their primary functions are to

The best characterised adult stem cells are those found in the bone marrow, which provides a unique niche for haematopoietic stem cells (HSCs) and the mesenchymal stem or stromal cells (MSCs). HSCs are responsible for the production and replacement of all blood cells during the entire lifetime 13. The earliest haematopoietic precursor, the haemangioblast, is not only a precursor of haematopoietic cell lineages but also of cells that

Mesenchymal stromal cells are a heterogenous population of stem/progenitor cells able to differentiate into several cell types such as chondrocytes, osteocytes, fibroblasts, myocytes, adipocytes, epithelial and neuron-like cells. When stimulated by specific signals, these cells can be released from their niche in the bone marrow into circulation and recruited to the target tissues where they undergo *in situ* differentiation and contribute to tissue homeostasis and repair 15. MSCs also secrete factors that promote survival and differentiation of endogenous cells as well as angiogenetic factors essential for blood vessel formation. MSCs possess remarkable immunosuppressive properties and can inhibit the proliferation and function of the major immune cell population 16, 17 as well as antimicrobial properties 18. Furthermore, these multipotential stromal stem and progenitor cells at different stages of maturation contribute to the formation of HSC stem cell niche and play a critical role in haematopoiesis 19.The characteristic of MSCs makes these cells exceptionally suitable for various therapeutic possibilities such as supporting tissue regeneration, correcting inherited disorders, dampening chronic inflammation, and Cellulose sponges can be manufactured by adding supportive strengthening fibres (8-10 mm long cotton fibres; about 20% of the weight of the cellulose) and sodium sulphate crystals as pore forming material to a cellulose viscose (sodium xantogenate) solution (4-6 g cellulose/100 g viscose). The cellulose is regenerated by heating the solution in a water bath after which the sponge is washed with hot water, treated with a dilute acid and sodium hypochlorite bleaching solution, and finally washed repetitively in distilled water before drying and sterilisation 20, 21. When inserted subcutaneously, a vital and well vascularised repair tissue, called granulation tissue, grows rapidly into this cellulose sponge. Due to this good granulation tissue formation ability, cellulose sponges have been used in experimental surgery for decades 6, 7, 22 and the subcutaneous implantation of the cellulose sponge is widely accepted method for wound healing (see box 2) studies 8, 23. Several cellulose products for wound healing purposes (e.g. Cellospon®, Cellstick®, Sponcal®, Visella®, and Absorpal®) are commercially available. These products are made from the sponge form viscose cellulose and have homogenous porous structure, characterized by thin pore walls with one or more inter-pore openings. They are elastic and can be compressed and expanded repeatedly with no damage to their internal structure, hence providing free entrance for the invading cells to the inner parts of the sponge 24.

Host reactions following implantation of biomaterials include injury, blood-material interactions, provisional matrix formation, inflammation, granulation tissue development, foreign body reaction, and fibrosis/fibrous capsule development 25. When implanted subcutaneously, a blood-material interaction occurs with protein adsorption to the cellulose sponge and a blood-based transient provisional matrix, a blood clot; is formed on and around the sponge. The platelets, originated from the injured blood vessels, not only participate to haemostasis but also liberate bioactive agents like cytokines and growth factors that will attract inflammatory and phagocytosing cells. The first cells to arrive are polymorphonuclear leucocytes, i.e. neutrophils, which are characteristic for the acute inflammatory response. These cells secrete pro-inflammatory cytokines that, in turn, attract circulating monocytes, which are activated and converted in the tissue to macrophages that kill bacterial pathogens, scavenge tissue debris and destroy remaining neutrophils. Biomaterial surface adherent macrophages can also fuse to form multinucleated foreign body giant cells. In their attempt to phagocytose the biomaterial, adherent macrophages become active 25. By releasing a variety of chemotactic, neovasculogenic and growth factors that stimulate cell migration, proliferation and formation of new blood vessels and tissue matrix, macrophages mediate the transition from the inflammatory phase to the

proliferative phase. During the proliferative phase, the provisional extracellular matrix in the cellulose sponge is gradually replaced with granulation tissue, which is formed from infiltrated mature fibroblasts and rapidly proliferating mesenchymal stromal cells (MSCs) differentiating to fibroblasts *in situ*. The newly formed extracellular matrix is rich in blood vessels, which carry oxygen and nutrients to maintain the metabolic processes. The sponge is surrounded by a well-vascularised fibrous capsule, which becomes somewhat thinner during the final remodelling phase [38].

Cellulose - A Biomaterial with Cell-Guiding Property 87

**Figure 1.** SEM micrograph of regenerated uncoated and HA-coated cellulose sponges (bar = 50 m). The hydroxyapatite layer was initiated in sterile 1 x SBF with bioactive glass at 37 °C for 24 h and was then grown in sterile 1.5 x SBF at the same temperature for 14 days under continuous shaking.

**Figure 2.** HA-coating of cellulose prevents bone in-growth**.** One year after implantation into rat femoral bone defect, new bone (nb) growth is mainly observed beneath (arrows) the HA-implant (a), which has been pushed out from the defect area. The HA-implant itself (b) is mostly filled with soft connective tissue containing abundant giant cells (arrow heads). Uncoated cellulose implant (c) allows new bone in-growth and the non-ossified parts contain less inflammatory cells. (a and c; van Gieson stain; b and d

haematoxylin-eosin stain; cf = cellulose fragment; scale bar = 100 m, modified from 11).

Similar biocompatible regenerated cellulose developed for wound healing studies has also been tested as a scaffold for cartilage tissue engineering. Although the cellulose sponge provided a non-toxic environment for cartilage cells, the construct remained soft and lacked the extracellular matrix composition typical for normal articular cartilage 26. When implanted into bone defects, regenerated cellulose strengthened by cotton fibres allowed new bone in-growth to some extent 9-11.

## *2.1.2. Hydroxyapatite-coating of regenerated cellulose*

The number of cells and tissue in-growth are affected to a certain limit by the porosity, size of pores, and the thickness of the pore walls of the cellulose sponge 8. We hypothesised that coating the regenerated cellulose with hydroxyapatite (HA) that resembles the mineral composition of bone, would improve its bone forming properties. The mineral originated from a specific bioactive glass, S53P4 (23% Na2O, 20% CaO, 4% P2O5, 53% SiO2) that has a good osteoconductivity and is in clinical use 27-32. However, glass as such, is difficult to trim to the desired size and form. Furthermore, it is brittle and fragile, and therefore, not suited in sites subjected to load like in femoral and tibial bone defects.

In our studies, the calcium phosphate layer was precipitated on cellulose sponges (10 x 100 x 100 mm) with average pore sizes between 50 and 350 m by the biomimetic method of Kokubo et al 33. Mineralisation was initiated in 500 ml of sterile simulated body fluid (SBF) supplemented with a 2.0 g of the bioactive glass at 37C for 24h and was then grown in 500 ml sterile 1.5 x SBF for 14 days at the same temperature under continuous shaking. The SBF solution was changed every second day. The formed calcium phosphate layer rich in silica was verified by scanning electron microscope (figure 1) and characterised with Fourier transform infrared spectroscopy 11. (1 x SBF = 136.8 mM NaCl, 4.2 mM NaHCO3, 3.0 mM KCl, 1.0 mM K2HPO4 x 3H2O, 1.5 mM MgCl2 x 6H2O, 2.5 mM CaCl2 and 0.5 mM Na2SO4, pH 7.4; ion concentration close to that of human plasma)

Sterile HA-cellulose and untreated cellulose sponges, sized 2.3 x 3 x 8 mm, were implanted into femoral bone defects of male rats aged 10-13 weeks (for further details see 11) and were followed up for 52 weeks. The implants were analysed histologically and with biochemical and molecular biologic methods. The HA layer did not improve the bone ingrowth into the cellulose sponge. In fact, the new bone was instead mainly formed beneath the implant at the bottom of the defect leaving the implant filled with a well vascularised fibrous tissue rich in inflammatory cells (figure 2). The inflammatory reaction was much stronger than in the uncoated cellulose indicated by the larger number of inflammatory cells

during the final remodelling phase [38].

new bone in-growth to some extent 9-11.

*2.1.2. Hydroxyapatite-coating of regenerated cellulose* 

7.4; ion concentration close to that of human plasma)

suited in sites subjected to load like in femoral and tibial bone defects.

proliferative phase. During the proliferative phase, the provisional extracellular matrix in the cellulose sponge is gradually replaced with granulation tissue, which is formed from infiltrated mature fibroblasts and rapidly proliferating mesenchymal stromal cells (MSCs) differentiating to fibroblasts *in situ*. The newly formed extracellular matrix is rich in blood vessels, which carry oxygen and nutrients to maintain the metabolic processes. The sponge is surrounded by a well-vascularised fibrous capsule, which becomes somewhat thinner

Similar biocompatible regenerated cellulose developed for wound healing studies has also been tested as a scaffold for cartilage tissue engineering. Although the cellulose sponge provided a non-toxic environment for cartilage cells, the construct remained soft and lacked the extracellular matrix composition typical for normal articular cartilage 26. When implanted into bone defects, regenerated cellulose strengthened by cotton fibres allowed

The number of cells and tissue in-growth are affected to a certain limit by the porosity, size of pores, and the thickness of the pore walls of the cellulose sponge 8. We hypothesised that coating the regenerated cellulose with hydroxyapatite (HA) that resembles the mineral composition of bone, would improve its bone forming properties. The mineral originated from a specific bioactive glass, S53P4 (23% Na2O, 20% CaO, 4% P2O5, 53% SiO2) that has a good osteoconductivity and is in clinical use 27-32. However, glass as such, is difficult to trim to the desired size and form. Furthermore, it is brittle and fragile, and therefore, not

In our studies, the calcium phosphate layer was precipitated on cellulose sponges (10 x 100 x 100 mm) with average pore sizes between 50 and 350 m by the biomimetic method of Kokubo et al 33. Mineralisation was initiated in 500 ml of sterile simulated body fluid (SBF) supplemented with a 2.0 g of the bioactive glass at 37C for 24h and was then grown in 500 ml sterile 1.5 x SBF for 14 days at the same temperature under continuous shaking. The SBF solution was changed every second day. The formed calcium phosphate layer rich in silica was verified by scanning electron microscope (figure 1) and characterised with Fourier transform infrared spectroscopy 11. (1 x SBF = 136.8 mM NaCl, 4.2 mM NaHCO3, 3.0 mM KCl, 1.0 mM K2HPO4 x 3H2O, 1.5 mM MgCl2 x 6H2O, 2.5 mM CaCl2 and 0.5 mM Na2SO4, pH

Sterile HA-cellulose and untreated cellulose sponges, sized 2.3 x 3 x 8 mm, were implanted into femoral bone defects of male rats aged 10-13 weeks (for further details see 11) and were followed up for 52 weeks. The implants were analysed histologically and with biochemical and molecular biologic methods. The HA layer did not improve the bone ingrowth into the cellulose sponge. In fact, the new bone was instead mainly formed beneath the implant at the bottom of the defect leaving the implant filled with a well vascularised fibrous tissue rich in inflammatory cells (figure 2). The inflammatory reaction was much stronger than in the uncoated cellulose indicated by the larger number of inflammatory cells **Figure 1.** SEM micrograph of regenerated uncoated and HA-coated cellulose sponges (bar = 50 m). The hydroxyapatite layer was initiated in sterile 1 x SBF with bioactive glass at 37 °C for 24 h and was then grown in sterile 1.5 x SBF at the same temperature for 14 days under continuous shaking.

**Figure 2.** HA-coating of cellulose prevents bone in-growth**.** One year after implantation into rat femoral bone defect, new bone (nb) growth is mainly observed beneath (arrows) the HA-implant (a), which has been pushed out from the defect area. The HA-implant itself (b) is mostly filled with soft connective tissue containing abundant giant cells (arrow heads). Uncoated cellulose implant (c) allows new bone in-growth and the non-ossified parts contain less inflammatory cells. (a and c; van Gieson stain; b and d haematoxylin-eosin stain; cf = cellulose fragment; scale bar = 100 m, modified from 11).

including macrophages and foreign body cells, which also is a sign of chronic inflammation. Activated inflammatory cells produce many pro-inflammatory bioactive agents, such as tumour necrosis factor-alpha (TNF-which is known to interfere with the bone specific transcription factor Cbfa1 and to depress the function of differentiated osteoblasts 34,35. Continuous exposure to these agents may, thus, inhibit differentiation of the progenitor cells into bone forming osteoblasts explaining, as least partly, the less osteoid tissue in HA-coated cellulose implants. Furthermore, the HA layer did intensify the attachment of transforming growth factor beta 1 (TGF) 11, a growth factor involved in fibroplasia. Hence, the HA surface did not offer any advantages in comparison with untreated cellulose in cortical bone defect healing.

Cellulose - A Biomaterial with Cell-Guiding Property 89

**Figure 3.** Oxidations with periodate and H2O2 increase the biocompability and degradation of cellulose. Oxidised cellulose (a, b) allows new bone (nb) formation when implanted into femoral bone (fb) defects of rat. (cs = cellulose scaffold, bm = bone marrow, m=muscle overlaying the implant site, arrow heads point at osteoblasts lining the new bone; haematoxylin-eosin stain; scale bars = 100 m (a), and 25 m

Biodegradation of cellulose can also be improved by treating it with urea. The resultant carbamino cellulose showed increased solubility that could be regulated by the duration of treatment. The fundamental aim was to develop material that could be used as a vehicle for drugs in tablets, or perhaps for subcutaneous long-lasting administration of drugs. Small, round or oval cellulose pearls with 50-500 m diameters can be manufactured from regular or carbamino cellulose by dropping viscose into a solution containing 100 g H2SO4 and 200 g Na2SO4/l at 20°C followed by centrifugation 36. Four and six per cent viscose solutions were used to make the 0.5 mm diameter cellulose pearls. The material was collected, washed with distilled water and 5g H2SO4/l and dried for 24 hours at 40°C. Sterilisation was carried

For implantation studies, several pearls were glued together with alginate 37 in moulds. The results from the subcutaneous implantation experiment (Figure 4 a, b) were encouraging as implanted 4% cellulose pearls were infiltrated with new granulation tissue and most of the pearls showed signs of nearly complete degradation where as 6%-pearls were more resistant during the observation period of two weeks. Intramedullary implantation into rat femoral bone (figure 3 c-e) showed similar behaviour: many of the 4 % pearls were infiltrated by new granulation tissue and some were surrounded by new osteoid tissue. There was some variation in the degree of degradation; while some pearls had been digested completely, some remained almost intact. No foreign body giant cells were observed, however. We do not know whether alginate surroundings affected the degradation of pearls in the bony environment, but to make the carbamino cellulose more useful in medical applications, the structure should be further altered to become even more vulnerable to hydrolytic enzyme attacks, especially if used for subcutaneous administration

(b)).

of drugs.

out by autoclaving or with 70% ethanol.

## **2.2. The effect of increased biodegradability of cellulose**

Another approach to improve the biocompatibility of cellulose was to alter its chemical structure in order to increase its biodegradability. The mild bleaching and oxidation of regenerated cellulose with sodium hypochlorite carried out during the preparation of cellulose sponge does not cleave the glucose ring and the resultant cellulose is not biodegradable, which probably prevented complete ossification of the implanted sponge. Therefore, in the search for suitable bone defect fillers, we extended the material development with a two sequential oxidation steps. Firstly, the cellulose was oxidated by periodate for 1-3 hours. This treatment opens some glucose molecules and should theoretically make them more susceptible to glucosidases and other enzymes capable for carbohydrate degradation. Excess periodate was washed by ascorbate or thiosulphate and water before the second oxidation by hydrogen peroxide (H2O2) for 3 or 4 hours. As the oxidation reactions were not complete, the resultant materials are combinations of 2,3 dialdehyde and 2,3-dicarboxyl celluloses. The biogradability of the celluloses was tested in SBF for 7, 15 and 30 days. Oxidations for 3 h in periodate followed by 4 h in H2O2 turned out to be the best combination as 70% of the material was dissolved. Therefore, this material was used for further testing. No cytotoxicity was observed in fibroblast cultures. The material has to be sterilised by 70-95 % ethanol or ethylene oxide because autoclaving destroys the porous structure of the scaffolds.

The results of the bone implantation experiments (figure 3 a, b) showed that oxidised scaffolds were flattened, their pores had disappeared and the material was completely replaced by cells so that no visible cellulose fibrils were observed in the implantation sites. The degradation was not complete as the phagocytosing cells were full of homogenous material. It is conspicuous, however, that no giant cells were observed in the oxidised samples, whereas normal cellulose always induces a number of foreign body giant cells. If the sponges were oxidised more extensively their structures collapsed and the material could not be used for implantation. The implanted scaffolds did not show, on the other hand, any significant bone in-growth. Instead they consisted of cell masses that histologically were strikingly homogeneous. New bone had been grown on the opposite site of the implant strengthening the defect site. Despite improved biodegradability, oxidised cellulose was considered to have no value as a bone substitute.

**2.2. The effect of increased biodegradability of cellulose** 

cellulose was considered to have no value as a bone substitute.

defect healing.

porous structure of the scaffolds.

including macrophages and foreign body cells, which also is a sign of chronic inflammation. Activated inflammatory cells produce many pro-inflammatory bioactive agents, such as tumour necrosis factor-alpha (TNF-which is known to interfere with the bone specific transcription factor Cbfa1 and to depress the function of differentiated osteoblasts 34,35. Continuous exposure to these agents may, thus, inhibit differentiation of the progenitor cells into bone forming osteoblasts explaining, as least partly, the less osteoid tissue in HA-coated cellulose implants. Furthermore, the HA layer did intensify the attachment of transforming growth factor beta 1 (TGF) 11, a growth factor involved in fibroplasia. Hence, the HA surface did not offer any advantages in comparison with untreated cellulose in cortical bone

Another approach to improve the biocompatibility of cellulose was to alter its chemical structure in order to increase its biodegradability. The mild bleaching and oxidation of regenerated cellulose with sodium hypochlorite carried out during the preparation of cellulose sponge does not cleave the glucose ring and the resultant cellulose is not biodegradable, which probably prevented complete ossification of the implanted sponge. Therefore, in the search for suitable bone defect fillers, we extended the material development with a two sequential oxidation steps. Firstly, the cellulose was oxidated by periodate for 1-3 hours. This treatment opens some glucose molecules and should theoretically make them more susceptible to glucosidases and other enzymes capable for carbohydrate degradation. Excess periodate was washed by ascorbate or thiosulphate and water before the second oxidation by hydrogen peroxide (H2O2) for 3 or 4 hours. As the oxidation reactions were not complete, the resultant materials are combinations of 2,3 dialdehyde and 2,3-dicarboxyl celluloses. The biogradability of the celluloses was tested in SBF for 7, 15 and 30 days. Oxidations for 3 h in periodate followed by 4 h in H2O2 turned out to be the best combination as 70% of the material was dissolved. Therefore, this material was used for further testing. No cytotoxicity was observed in fibroblast cultures. The material has to be sterilised by 70-95 % ethanol or ethylene oxide because autoclaving destroys the

The results of the bone implantation experiments (figure 3 a, b) showed that oxidised scaffolds were flattened, their pores had disappeared and the material was completely replaced by cells so that no visible cellulose fibrils were observed in the implantation sites. The degradation was not complete as the phagocytosing cells were full of homogenous material. It is conspicuous, however, that no giant cells were observed in the oxidised samples, whereas normal cellulose always induces a number of foreign body giant cells. If the sponges were oxidised more extensively their structures collapsed and the material could not be used for implantation. The implanted scaffolds did not show, on the other hand, any significant bone in-growth. Instead they consisted of cell masses that histologically were strikingly homogeneous. New bone had been grown on the opposite site of the implant strengthening the defect site. Despite improved biodegradability, oxidised

**Figure 3.** Oxidations with periodate and H2O2 increase the biocompability and degradation of cellulose. Oxidised cellulose (a, b) allows new bone (nb) formation when implanted into femoral bone (fb) defects of rat. (cs = cellulose scaffold, bm = bone marrow, m=muscle overlaying the implant site, arrow heads point at osteoblasts lining the new bone; haematoxylin-eosin stain; scale bars = 100 m (a), and 25 m (b)).

Biodegradation of cellulose can also be improved by treating it with urea. The resultant carbamino cellulose showed increased solubility that could be regulated by the duration of treatment. The fundamental aim was to develop material that could be used as a vehicle for drugs in tablets, or perhaps for subcutaneous long-lasting administration of drugs. Small, round or oval cellulose pearls with 50-500 m diameters can be manufactured from regular or carbamino cellulose by dropping viscose into a solution containing 100 g H2SO4 and 200 g Na2SO4/l at 20°C followed by centrifugation 36. Four and six per cent viscose solutions were used to make the 0.5 mm diameter cellulose pearls. The material was collected, washed with distilled water and 5g H2SO4/l and dried for 24 hours at 40°C. Sterilisation was carried out by autoclaving or with 70% ethanol.

For implantation studies, several pearls were glued together with alginate 37 in moulds. The results from the subcutaneous implantation experiment (Figure 4 a, b) were encouraging as implanted 4% cellulose pearls were infiltrated with new granulation tissue and most of the pearls showed signs of nearly complete degradation where as 6%-pearls were more resistant during the observation period of two weeks. Intramedullary implantation into rat femoral bone (figure 3 c-e) showed similar behaviour: many of the 4 % pearls were infiltrated by new granulation tissue and some were surrounded by new osteoid tissue. There was some variation in the degree of degradation; while some pearls had been digested completely, some remained almost intact. No foreign body giant cells were observed, however. We do not know whether alginate surroundings affected the degradation of pearls in the bony environment, but to make the carbamino cellulose more useful in medical applications, the structure should be further altered to become even more vulnerable to hydrolytic enzyme attacks, especially if used for subcutaneous administration of drugs.

Cellulose - A Biomaterial with Cell-Guiding Property 91

**Figure 5.** a). A schematic presentation of the subcutaneous implantation model used in our studies. Two midline incisions were made on the back of the rats, and sterilised, moistened sponge implants (10

Subcutaneously implanted cellulose sponges 7 days after implantation. HA-coated implants are darker in colour as a sign of high cellularity and rich neovascularisation, whereas the uncoated implants are

**Figure 6.** The HA-coating accelerated tissue growth into subcutaneously implanted cellulose sponges as well as the inflammatory response and blood vessel formation. a) Haematoxylin-eosin–stained sections 1 (upper), 3 (middle), and 7 (lower) days after implantation. The arrows in HA-coated sponges point at the border between the implant and the surrounding capsule (scale bar = 100 m). b) HAcoated sponges contain large clusters (arrows) of accumulated macrophages (brownish coloured cells). Macrophages favour gathering near to cellulose fibres (arrow head) (day 5; scale bar = 50 m). c) More blood vessels, as indicated by CD31-staining, can bee seen in 5-day-old HA-coated sponge compared to

x 5 mm) were inserted bilaterally into subcutaneous pockets under general anaesthesia. b).

pale.

uncoated one (scale bar = 50 m).

**Figure 4.** Tissue reactions of carbamino cellulose two weeks after implantation. Subcutaneously implanted 6%-cellulose pearls (p) stayed intact and showed only modest degradation (a), whereas b) 4%-cellulose pearls were degraded and infiltrated with new granulation tissue (gf). Similar behaviour was observed in bone implants: c) 6%-cellulose pearls were surrounded by a thin connective tissue capsule (arrow) whereas about half of the b) 4%-cellulose pearls were partially degraded and surrounded by bone (nb) or a thin osteoid layer (ol) even in the bone marrow (bm) area. (van Gieson stain; equal magnifications; scale bar 200 m).
