*3.3. Thermogravimetric Analysis*

The thermal degradation behaviors and DTG curves of Entry 12C are shown in Figure 7, and some crucial data are also displayed. The degradation curve of the crosslinked polyester in the first step was almost the same as the pure polyester, which was reflected by the same 5% weight loss temperature (327 ◦C) and the maximum weight loss rate temperature (403 ◦C). This step degradation was attributed to the ester and aliphatic moieties. The Tmax2 (455 ◦C) was observed in the cross-linked sample, which indicated the cross-linked polyester degraded by a multistep process. The reason might be that the hard segment urethane bonds formed after cross-linking, which reduced the migration of the chain segments. The thermal stability of the material was enhanced after cross-linking [61].

**Figure 7.** TG and DTG of NWCPE and cured polyester film.

#### *3.4. Tensile Properties of the Cross-Linked NWCPE Films*

The tensile properties of cross-linked NWCPE films were summarized in Table 5, and the stress–strain cures were listed in Figure 8. Mechanical properties are crucial for polymer materials which directly affect their industrial applications. The tensile strength increased from 3.1 MPa to 5.9 MPa after CHDM being introduced into the molecule. But when the content of CHDM reached 40%, the strength of the polymer no longer increased with the increase in the CHDM content. In addition, the toughness even decreased slightly due to the excessively high content of this rigid monomer. The result may be related to the following factors. First, the utilization of CHDM enhanced the rigidity of the polymer which could restrict the movement of the molecular chains, and the intermolecular force was increased, resulting in improved mechanical properties. Secondly, the molecular weight of the polymers were slightly different. Polyester with minor *M*n contained more terminal hydroxyl groups in the same mass, resulting in a slightly larger cross-linking density during the curing process. The stress–strain curves of all of the films showed typical plastic deformation, and the elongation at the break (ε) of the films were all basically above 70%. The cause might be the following two factors. First, the synthesized polyester has low crystallinity, leading to high mobility between molecular chains. Secondly, the polyetheramine side chains are similar to a spring and increased the ductility of the molecular. The tensile properties of cured NWCPE were better than some cured ionic waterborne polyesters, some of which were about 2–3 MPa [3,62,63].

**Table 5.** Tensile properties of cross-linked films.

**Figure 8.** Stress-strain profiles of cross-linked films.

#### *3.5. The Morphology and Light Transmittance of the Cured Film*

As shown in Figure 9a,c, the neat cross-linked film was transparent, and, at the entire visible wavelength range, the light transmittance of the film was high. In addition, owing to the absence of emulsifier, no surfactant molecules migration during the filmforming process was observed, and the surface of the film was smooth and flat according to Figure 9b [64]. At the same cross-linking density, the CHDM content exerted almost no influence on the transmittance of the film. The SEM and photographic images of cured

Entry 12C (CHDM 40%) were taken as examples. The high transparency was attributed to the following reasons. First, the synthesized NWCPEs have good compatibility with the curing agent WHDIT and there was not significant macrophase separation during the cross-linking procedure. Second, the surface of the film was smooth, which was hard for scattering light [65]. As such, the low crystallinity of the WCPE may also contribute to the high transmittance of the film.

**Figure 9.** The light transmittance (**a**), SEM image (**b**), and photograph (**c**) of cross-linked films.

### **4. Conclusions**

In this paper, MA, HA, CHDM, HG, and NPG were first used as monomers to synthesize linear polyesters by step-growth polycondensation. After this, the hydrophilic segment Jeffamine M-1000 was grafted to the polymer backbone by aza-Michael addition to prepare the NWCPE. The optimal reaction conditions were verified as TsOH as the selected catalyst, a reaction temperature of 180 ◦C, a MA/HA ratio of 5:5, and the content of the rigid monomer CHDM at 40%. In this condition, NWCPE with a molecular weight of about 8000 g mol−1 was prepared. After dispersing into water, the dispersion with a solid content of 40% and the particle size was appropriate with narrow distribution. The dispersion was good in storage stability and remained stable after 6 months. The good hydrolysis resistance was identified by the slight drop in pH after two months and little change in *M*n of the dispersion. After cross-linking with the curing agent, a transparent film with good mechanical properties and water resistance was obtained. The hardness of the film was H, and the adhesion level on the tinplate reaches level 0. The tensile strength of the film was 5.9 MPa, and the *ε* was about 88%. This work provides ideas for the synthesis and functionalization of waterborne polyesters. These synthesized polymers have wide application prospects in environmentally friendly polymer materials.

**Author Contributions:** Conceptualization, methodology, and validation, H.F., L.G. and S.G.; writing—original draft preparation, H.F.; writing—review and editing, H.F., L.G., and S.G.; supervision, and funding acquisition, L.G. and S.G. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research received no external funding.

**Institutional Review Board Statement:** Not applicable.

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** Not applicable.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**


**2.5 Screening and confirmation of ESBL and integrons producers**

A double synergy test was used for ESBL-producing strains testing. This consisted of placing disks (2–3 cm diameter) of ceftriaxone and cefotaxime around an amoxicillin-clavulanic acid disk on the bacterial plate.

For molecular characterization, DNA extraction was performed using heating method [23]. A loopful of bacterial growth from Mueller-Hinton agar (Liofilchem, Italy) plate was suspended in 1 ml of sterilized water. The mixture was boiled for 10 min at +100°C and centrifuged for 10 min at 12000 rpm at +4°C. Supernatant was then collected and used for the PCR reactions as DNA matrices. Multiplex PCR assays were performed for detecting EBLS-encoding genes (*bla*TEM, *bla*SHV, *bla*OXA, and *bla*CTX-M) and the presence of the class 1, class 2, and class 3 integrons from the β-lactams-resistant DEC strains. Primers (GeneCust, France) used for these amplifications are described in **Table 1**.

Thermocycling conditions were as follows: 5 min at +94°C, followed by 35 amplification cycles of +94°C for 30 s, 59 ± 4°C for 60 s, and +72°C for 60 s with a final extension of +72°C for 10 min on a thermal cycler (Gene Amp 9700, Applied Biosystems). PCR products were revealed on 1.5% stained Redsaf agarose gel (Prolabo, France), after electrophoresis under UV light (Gel Logic 200).

The PCR assays were carried out in a 25-ml reaction mixture, which consisted of 2.5 μl of the supernatant added to 22.5 μl of reaction mixture. This mixture contained 5 U of Taq DNA polymerase (Accu Power, South Korea), deoxyribonucleic triphosphate (10 mM), buffer GC (10X), MgCl2 (25 mM), and PCR primers (10 μM). Thermocycling conditions were as follows: 5 min at +94°C, followed by 35 amplification cycles at +94°C for 30 s, + 59 ± 4°C for 60 s, and +72°C for 60 s with a final extension of +72°C for 10 min on a thermal cycler (AB Applied Biosystems). Following PCR, the reaction products were separated using electrophoresis in 1.5% agarose gel (weight/volume), stained with Redsaf solution (Prolabo, France), and visualized under UV light (Gel Logic 200) [23].

*Beta-Lactamase-Producing Genes and Integrons in* Escherichia coli *from Diarrheal Children… DOI: http://dx.doi.org/10.5772/intechopen.103169*

**Table 1.**

*List of all primers used for antibiotic ESBL genes and integrons detection.*


**2. Bone Tissue**

Bone tissue is a diverse form of connective tissue with high metabolic activity, heterogeneous and dynamic structure and high mechanical strength [18,19]. It is made of extracellular substances and bone cells: osteoblasts, osteoclasts, osteocytes and osteogenic cells. Bone tissue co-creates the locomotor system, protects internal organs and bone marrow and stores mineral salts (99% calcium, 88% phosphorus, 50% magnesium and 35% sodium are located in bone tissue) [19]. Adaptation of bone structure to perform such important functions includes a number of organizational levels. These include: the molecular structure and distribution of crystals and organic components (nanoscale); the structure of bone plates; the structure and arrangement of spongy bone tissue and osteons of compact bone tissue; and macroscopic structure (macroscale) (Figure 1) [20,21].

**Figure 1.** The multi-scale structure of natural bone.

The bone has a hierarchical structure: from the level of the whole tissue, i.e., the occurrence of various types of long and short bones, flat or tubular, to the level of the tissues which are arranged in cortical and spongy structures, through to the microscopic level (images of cells, matrices and minerals) to the level of nanometers from single bone apatite crystals and collagen fibers [20,22].

The osteon is the main structural and functional unit of compact bone. Its structure is made of 6–15 cylindrical bone plates arranged concentrically around the central channel (Haversian canal). The interior of the canal is filled with individual osteogenic cells, osteoblasts and osteoclasts. The diameter of the channels is 20–100 μm. Bone plates are made of parallel fibers, mainly built with collagen type I mineralized with nanocrystalline multisubstituted carbonate hydroxyapatite (so-called bone apatite). Bone tissue also includes other non-collagen proteins and water. In general, apatite ensures bone hardness, while the organic fraction forms the scaffold for the biomineral and regulates the biomineralization process [21].

Bone apatite, i.e., biological apatite, is a mineral with a specific chemical composition that determines the biological, physicochemical and mechanical properties of the entire tissue. It is a nanocrystalline carbonate hydroxyapatite, additionally containing a variety of different ions (e.g., Mg2+, K+, Na+, Mn2+, HPO42− and SiO44−) [23–26]. The organic fraction of bone tissue is mainly made up of type I collagen [27]. The latter consists of three polypeptide chains entwined to form a triple helix. This structure is a so-called superhelix, with the occurrence of characteristic fragments containing repeating sequences: Gly–Pro–Hyp. Five triple helices assemble with each other, creating a microfibril. The microfibrils then organize into fibrils, forming compact fibers with diameters of about 100–200 nm. Collagen fibers also crosslink via lysine residues. The ordered fibrillar system is stabilized by other non-collagen proteins [27]. In the free spaces of collagen fibers, apatite crystals with a width of 15–30 nm, a length of 30–50 nm and a thickness of 2–10 nm are settled. They are composed of a crystalline core and a hydrated surface layer. The hydrated surface layer is about 1–2 nm thick and contains various ions. Recent studies show that the components of the hydrated surface layer are responsible for apatite reactivity, adsorption properties and the process of crystal maturation and growth. It is noteworthy that this is also the border region between crystalline apatite and the organic matrix of bones [28–31]. The small sizes of the crystals and the presence of such ions as CO32−, Na+ and Mg2+ (in amounts of approximately 4–6%, 0.9% and 0.5%, respectively) means that particles of natural hydroxyapatite are easily absorbed [32].

Among non-collagen proteins of extracellular bone tissue (constituting about 5% of the organic matrix and synthesized by osteoblasts or other cells, or reaching bone tissue with blood), we distinguish proteoglycans (mainly chondroitin sulphate), which can affect the formation, thickness and orientation of collagen fibers, and are also for the binding of hydroxyapatite, osteonectin and osteocalcin, and fixing proteins such as fibronectin, osteopontin and bone sialoprotein (BSP) [18,19].

In addition to the organic matrix (approximately 20–30% by weight) and the mineral fraction (representing 60–70% of bone mass and consisting primarily of nanocrystalline apatite), the third equally important component of bone tissue is water. It constitutes about 10% of bone mass. It facilitates fluid transport, contributes to elastic properties and plays a key role in the mineralization process. Water is on the surface of mineral crystals, inside the crystals and between collagen fibers. Most of the water is located in the pore spaces (so-called associated water), whose content decreases with bone age. Apart from water associated with bone tissue, there is also bound water, located in the organic matrix; i.e., in type I collagen fibers and in bone mineral.

It is also worth noting that bone tissue is a dynamic structure whose composition is not constant. The content of water in bone tissue is also variable. This can affect the properties of collagen (its elasticity); the coherence of the connections of bone composite components; and the ion exchange and balance between bone formation and bone resorption. This is why the formation of new biomimetic HA/Col biocomposites which imitate biological bone tissue is so important [23,27].

### **3. Synthesis of Collagen–Apatite Composites**

When producing HA/Col composites, one should aim to produce a material that closely resembles the chemical composition, micro and macro-structure and porosity of the natural composite, i.e., bone tissue. Thanks to these biomimetic properties, materials with high compatibility as well as

osteoinduction and ocostoconductivity can be obtained [4,33]. There are several methods for obtaining HA/Col composites [34].

Among them, the most basic method is the simple mixing of previously obtained apatite powder with a collagen solution. Apatite can be synthesized in many ways. There are many reviews on the preparation and properties of hydroxyapatite and hydroxyapatite enriched with various ions [35–38]. Usually, wet methods are used, which involve the precipitation of calcium phosphate from appropriate reagents (e.g., calcium nitrate and ammonium phosphate, as sources of calcium and phosphorus, respectively) added in the appropriate ratio (Ca/P molar ratio = 1.67) and at the appropriate pH (usually pH > 8). The precipitate, after an appropriate aging time, is subjected to filtration, drying and heating at a suitable temperature. It is worth noting that the concentrations of reagents, temperature, pH and aging time influence the size of the crystals obtained and their morphology [35,36].

Type I collagen is often used to make composite materials (this is the organic matrix of bone tissue). Collagen can be obtained from pig skin, bovine or horse tendons, rat tails, etc. [39]. In some works, atelo-collagen was used, which was obtained after the enzymatic treatment of collagen and removal of telopeptides (to minimize antigenecity) [40,41]. Atelo-collagen is often more soluble and forms a collagen solution, whereas type I collagen forms a suspension. It is also worth mentioning that collagen used as an individual material does not have osteoinductive properties, but acquires them in combination with calcium phosphate (apatite) [3,4]. After mixing the gel/collagen solution with apatite powder, a suspension is formed, which is then subjected to drying at a critical point, or lyophilization. Of course, many authors have applied modifications of this method [42–55].

An interesting comparison of two methods for the preparation of HA/Col composites was presented by Cuni ffe et al. [42]. In the first method, nanohydroxyapatite particles were added to the collagen suspension (slurry-suspension method) and then lyophilized, while in the second method, the lyophilized collagen in the form of a porous sca ffold was soaked in nHA suspension and then lyophilized (immersion method). In both cases, composite sca ffolds with highly porous, interconnected structures were obtained. It was found that the suspension method was more repeatable and easier to perform. In a paper by Uskokovi´c et al. [43], hydroxyapatite was obtained by the reaction of ammonium phosphate and calcium nitrate and then calcinated at 1100 ◦C during 6 h. The hydroxyapatite was mixed with type II collagen in a mortar and then pressed into pellets at room temperature and 60 ◦C. The resultant composite material was subjected to physicochemical analysis. SEM analysis showed that the material pressed at a higher temperature is characterized by more intimate contact between collagen type I and apatite phases.

In a paper by Cholas et al. [44], hydroxyapatite microspheres obtained by spray drying were used to produce the HA/Col hybrid composite. The paper suggests the possibility of using such a composite as a carrier for a drug substance that would be placed in the mesoporous structure of the microspheric HA (Figure 2).

Since 3D porous materials are characterized by a spongy structure, an interesting solution was proposed by Teng et al. [45]. Type I collagen dissolved in 1,1,1,3,3,3-hexafluoro-2-propanol (HFP) was added to the aqueous suspension of hydroxyapatite. Titanium discs after previous cleaning were coated with a homogeneous mass of a composite in a spinner, followed by drying the coated Ti substrates in a desiccator under vacuum overnight. Subsequently, the coatings were chemically cross-linked in two solutions: N-(3-dimethylaminopropyl)-N0-ethylcarbodiimide (EDC) hydrochloride and N-hydroxysuccinimide (NHS). The composite coatings obtained were characterized by high homogeneity, while the sample containing 20% hydroxyapatite turned out to be the most hydrophilic.

**Figure 2.** SEM of scaffolds. (**a**,**b**) Pure collagen. (**c**,**d**) Col/mHA (collagen/hydroxyapatite-microsphere). Scale bars: 100 μm (**a**,**c**), 20 μm (**b**,**d**). Reprinted from [44] with permission from Elsevier.

Tampieri et al. obtained (HA/Col) composites by two methods. In the first, a collagen suspension was mixed with hydroxyapatite previously obtained by precipitation from Ca(OH)2 solution with H3PO4 solution [46]. In the second method, the precipitation of hydroxyapatite from the reagents used in method I was carried out in the presence of collagen. From the results, it can be concluded that the method based on the direct precipitation of apatite in collagen solution is best. The resultant material has much greater similarity to bone tissue, and the collagen fibers are in close connection with apatite crystals. The material obtained by simply mixing the components is characterized by properties similar to those of collagen. In addition, it is worth noting that according to the SEM results, the composite obtained by the standard method (simple mixing hydroxyapatite with collagen) has a less homogeneous structure with apatite crystals distributed unevenly on the surface of collagen fibers.

The method with the precipitation of hydroxyapatite in the presence of collagen has been repeated by many researchers [30,56–60].

An interesting modification was developed by Yunoki et al. [56]. Synthesized under standard conditions (phosphoric acid and Ca(OH)2 as reagents for HA and atelo-collagen type I mixed together at pH 8–9 and temperature around 40 ◦C), self-organized nanocomposite HA/Col (after initial lyophilization) was placed in distilled water or in PBS buffer (standard PBS, pH = 7,4) and then re-frozen and placed under vacuum at 140 ◦C. In the presence of PBS, more effective crosslinking of collagen fibers occurred and porous composites with very good mechanical properties were obtained. According to Krishnakumar et al. [57], ribose can be used to successfully crosslink the HA/Col composite structure (Figure 3). In this work, MgCl2•6H2O was used as a source of magnesium ions introduced into the structure of synthesized apatite in order to obtain a highly compatible material with bone tissue. In research by Calabrese et al. [58], magnesium ions were also used in the production of the composite, but crosslinking was carried out using bis-epoxy (1,4-butanediol diglycidyl ether, BDDGE). Glutaraldehyde was another reagent that was used to crosslink the obtained hybrid composites [30,49].

**Figure 3.** Detailed schematic illustration of the MgHA/Coll (type I collagen matrix with magnesium-doped-hydroxyapatite nanophase) hybrid scaffold's development: (**A**) pH-driven, bioinspired biomineralization process; (**B**) MgHA/Coll crosslinking with ribose scaffolds in pre and post-glycation processes. Reprinted from [57] with permission from Elsevier.

A completely different approach to creating a composite 3D structure was presented by Zhou et al. [59]. To improve the mechanical properties of the composite, a porous ceramic matrix of hydroxyapatite and beta-tricalcium phosphate (beta-TCP) was first formed. Then the ceramic 3D material was soaked in a collagen and SBF suspension at pH 4–6 and then completely immersed in a collagen solution and closed under high pressure. A vacuum infusion was carried out at a pressure of 10 Pa and held for 2 h to allow complete saturation of the samples. The scaffolds were freeze-dried and then crosslinked with glutaraldehyde.

### **4. Biological Properties of Apatite–Collagen Composites**

Of course, a key aspect of research on HA/Col composites is obtaining information on their biological properties and the potential uses of the resulting biomaterials. Recently, there have been more studies on this topic. Preclinical and clinical studies using various in vitro and in vivo models provide important information on bone tissue regenerative properties; bioresorbability; and the impacts of the elasticity and porosity of such bone substitutes in their micro and macro-environments.

### *4.1. Osteoconductivity, Osseointegration, Bioactivity and Biocompatibility*

There are many studies confirming the ability of HA/Col composites to stimulate the formation of new bone tissue. In all biological tests for the production of HA/Col composites, type I collagen (atelo-collagen extracted from porcine dermis or bovine tendon) was used. In one study, Fukui et al. implanted composites consisting of nano-HA and collagen into the mandibles of rabbits [61]. Collagen sponge and collagen sponge/calcined hydroxyapatite composite were used as controls. Calcined HA was prepared by heating nano-HA at 900 ◦C for one hour. Histological examination showed a greater amount of newly formed bone tissue in the nano-HA/Col composite environment than in controls, and faster implant replacement with host bone tissue, which also confirms the high bioresorbability of the material.

The purpose of the research by Kikuchi et al. [40], was to synthesize a composite material (based on hydroxyapatite and collagen) as similar as possible to bone tissue and to use it for testing on dog tibia bones. In the synthesis of composites, the optimal pH and temperature were selected to improve the mechanism of collagen fiber organization. The composite was introduced in place of a 20 mm defect. Bone condition was observed using X-ray techniques for 12 weeks. In the place of the composite, the newly formed bone tissue gradually penetrated, and the defect was completely filled after 8 weeks. After 12 weeks, the bone and the composite were removed and the material was observed with a microscope (Figure 4). Two types of cells near the composite were observed— osteoblasts and osteoclasts. The composite was included in the bone remodeling process, which so far has mainly been observed for autografts. Such high bioactivity of the material may be the result of high similarity to bone tissue. The composite can be recognized as bone by the surrounding cells. Nishikawa conducted research on a similar group (dog tibia bones) [62]. The stimulating e ffect of such composites on the synthesis of new bone tissue has been proven. It is concluded from the results of the study that HA/Col may be a source of calcium ions that are incorporated into the newly formed bone tissue.

**Figure 4.** HE-stained histological section (×100) of HAp/Col composite implanted into a beagle's tibia for 12 weeks. Triangles indicate elongated cells, and arrowheads multinucleated giant cells. Reprinted from [40] with permission from Elsevier.

An important property when describing the biological functions of HA/Col composites, is their ability to support migration of cells from surrounding tissues. Yoshida et al. studied the adhesion, proliferation and osteogenic response of MG63 cells using 3D sponges, high porosity HA/Col and as a control, collagen sponge [63]. The cells with sponge were examined by histology, total DNA content and gene expression. The results suggest that materials based on hydroxyapatite and collagen have good osteogenic properties and can successfully serve as sca ffolds in bone tissue reconstruction. The total DNA content in the HA/Col sponge was 1.8 times higher than in the control sample, and the osteogenic cells showed good and even adhesion over the entire surface of the sponge. HA/Col composites create a space that facilitates precursor cell migration, proliferation and di fferentiation. The results were confirmed by another study using a HA/Col membrane [64]. Similarly, bone marrow cells were cultured together with osteoblasts on the HA/Col composite and di fferentiation to osteoclasts was observed without the addition of other factors. That distinguishes this material from pure HA or TCP [65]. Wu et al. prepared the HA/Col composite in the form of microspheres, and with this model also proved that osteoblasts are capable of proliferation, di fferentiation and mineralization in the matrix of microspheres (Figure 5) [47].

**Figure 5.** Confocal microscopy of osteoblast cells cultured on microsphere staining with DNA dye YOYO-1 (HAp: hydroxyapatite, OB: osteoblast; 4 days after seeding). Reprinted from [47] with permission from Elsevier.

Calabrese et al. implanted HA/Col composites in mice. Hydroxyapatite was additionally enriched with magnesium ions [58]. The results indicate the ability of this material to recruit host cells and promote ectopic bone growth *in vivo*. Correct angiogenesis was also confirmed by FMT analysis. The authors emphasize that the materials are characterized by a high degree of safety, due to the lack of the addition of growth factors or cells subjected to in vitro manipulation, and they can be a safe and promising solution in the treatment of bone diseases [66].

There is a special, commercially available bone substitute biomaterial (Biostat) with the composition: hydroxyapatite, collagen and chondroitin sulphate. In a clinical trial, two groups were followed: group A—filling the defect with Biostat; group B—no defect fill [67]. After 4–6 months, the group treated with Biostat implants showed a higher percentage of new bone tissue coverage (67%) compared to group B (34%). Another study confirms that Biostat material affects bone tissue reconstruction [68].

Materials with the best biocompatibility are still being sought. Certainly, the greater the similarity to physiological bone tissue, the more likely it is that tissue compatibility will be achieved. It is presumed that the carbonate content of apatite affects the formation of new bone tissue by affecting the solubility and crystallinity of apatite [69]. Biological apatite contains about 4–6 wt% of carbonates [24]. Matssura et al. proved that among the HA/Col composites with different carbonate contents, the one with the content of 4.8% could be distinguished from the others [48]. Composites containing this apatite had the greatest ability to form new bone tissue after implantation in the femurs of rabbits. This was observed on the X-rays and examined using histology. It was also concluded that bone metabolism is strongly associated with the physicochemical properties of apatites, especially with the carbonate content.

Mazzoni et al. assessed the biocompatibility, osteoconductive and osteoinductive properties of HA (Pro Osteon 200) and collagen (Avitene) composites using a cell model—mesenchymal stem cells (hMSC) [70]. Expression of osteogenic genes was analyzed in cells located on the composite. The results showed that such biomaterial has the ability to induce osteogenic differentiation of hMSC, because it induces osteogenic genes and increases matrix mineralization without toxic effects. Other studies also conducted on the cell line have confirmed that the HA/Col composite has osteoinductive properties and is a good tool to accelerate the migration, proliferation and differentiation of bone tissue cells [71]. The above material was also used in maxillofacial surgery as a kind of scaffolding for the zygomatic bone. The high biocompatibility and the osteoconductive properties of the composite have been confirmed and the low number of postoperative infections has been noted [72].

Initial tests have been carried out to check the bone density change after using various defect replacement materials. One study used computer tomography (CT) to evaluate extraction site

dimensions and density changes after a tooth extraction. Di fferent graft materials were tested [73]. Patients were divided into three groups. The first group was treated with a demineralized bone matrix with the addition of collagen membrane and the second with hydroxyapatite with the addition of collagen membrane; in the third group the extraction site remained empty. A CT scan was performed 10 and 120 days after surgery. It was shown that the use of HA/Col gives the highest bone density in CT compared to the use of demineralized bone matrix and group III. However, no changes in vertical socket dimension were observed. Still, the authors of the study say that HA and collagen-based material could be recommended to improve bone quality and could prepare the extraction site properly for proper implant placement.

Clinical studies have also compared HA/Col and β-TCP as implants in patients after a history of bone cancer or fractures [74]. The e ffectiveness of the materials was evaluated by using X-rays to assess bone regeneration. This study proved the superiority of porous HA/Col over β-TCP by presenting the results of bone regeneration and implant resorption (Figure 6). In contrast to β-TCP, the material containing collagen adapted to the shape of the bone defect, leaving no gaps or free spaces, and connected continuously with the bone. There were more side e ffects for HA/Col than for β-TCP, but they were not serious.

**Figure 6.** Results of X-ray evaluation. The scores improved over time during the follow-up period in both groups. At each time point, the score in the HAp/Col group was higher than that in the β-TCP group. Reprinted from [74] with permission from Elsevier.

### *4.2. Bioresorbability of Composites*

Biodegradation or bioresorbability are very desirable processes when designing implant biomaterials. Due to such properties, the implanted material quickly disappears and is replaced by newly forming host bone tissue. Biodegradation has been studied using composite membrane carbonate apatite–collagen, in vivo and in vitro [75]. With the in vitro method, the membranes were immersed in collagenase solution and the degradation time of the composite was analyzed. The results showed a gradual increase in the concentration of calcium ions, which is related to and dependent on the dissolution of the collagen membrane. The membrane was also implanted in rat bone tissue and histological changes were analyzed. Studies have shown good membrane biocompatibility and the biodegradation time has been reduced due to the presence of carbonate apatite. The biodegradation time can therefore be controlled by the carbonate apatite contained in the membrane.

Another study compared the properties of pure hydroxyapatite with a carbonate apatite–collagen composite [76]. Rabbit tibia bone fragments were removed and appropriate materials were implanted. Composites gradually degraded, and the newly formed bone filled the defect within 6 weeks, while

the implanted hydroxyapatite was not replaced by bone tissue. There are also other animal studies that confirm the successful osseointegration of HA/Col materials [61,77].

### *4.3. Porosity and Biological Properties*

The porosity of HA/Col composites also has a significant impact on biological properties. This was tested on an animal model by implanting porous scaffolds into the tibias of rabbits. Implants with higher porosity were characterized by faster formation of new bone tissue and its penetration into composite structures. The porous structure ensures better osteoconductivity. It allows better cell migration and facilitates the formation of blood vessels that ensure the proper nutrition of newly formed bone tissue [78–80]. Porous composites in contact with water become elastic. This property means that they are easy to use in surgery and adapt to the difficult shapes of cavities. However, it is believed that these materials are less mechanically strong due to their porous structures. In the study using HA/Col and TCP, the above materials were implanted in the tibia bones of rabbits to test their biomechanical properties. It was found that although the HA/Col composite is less mechanically durable than TCP, after implantation in the place of defect it showed greater mechanical strength in a given place than the TCP material [81].

Despite the fact that high porosity reduces its mechanical strength, sponge-like elasticity provides easy "handling" during surgery [82]. After wetting, HA/Col porous composites become elastic, like a sponge, and are easily implanted in bone defects [78,83]. From a surgical point of view, the addition of collagen to hydroxyapatite provides many beneficial features during surgery: ease of fitting to the defect site (blocks can be cut), easy adaptation to the defect morphology, the ability to stick material to the transplant site and the ability to promote clot formation and stabilization thanks to the hemostatic properties of collagen [63,79].


**Author details**

materials to be shipped greater distances for further processing into molecules that provide the greatest benefit either as biofuels or as precursors for other organic based materials.

One of the challenges of any biomass conversion platform is dealing with the fermentation residual materials. Lignin is a primary component of the fermentation waste and in many schemes it is recovered and burned to supply energy for other steps in the complete process. With the carboxylate platform based upon mixed ruminal microbes, one of the by products could be the microbial protein as a value-added material. In the normal rumination process, formation of microbial protein is an important component to supply needed protein to the animal. In dairy production, microbial protein helps supply critical amino acids required for milk production, especially methionine and lysine that are often low or lacking in many foragebased diets [52]. Harvesting the microbial protein after biomass conversion to biofuels could provide an important protein supplement for dairy cow diets that is enriched in methionine and lysine. The microbial proteins would be insoluble along with the typical insoluble materials, i.e., lignin and other cell wall components. Recovery of these insoluble materials would be relatively straightforward. As an alternative the lignin-microbial-carbohydrate residue from the fermentation process could be used to replace phenolic-formaldehyde based adhesives[53]. Many of the ruminal microbes contain glycocalyx materials surrounding the individual cells that help them adhere to plant materials during digestion. The glycocalyx is a glycoprotein-polysaccharide complex that surrounds the cell membrane of some bacteria[54]. It has also been demonstrated that the lignin-microbial residues from ruminal fermentations, as proposed for the carboxylate platform, could be used to replace phenol-formaldehyde compounds as adhesives in the production of plywood composites[53]. Up to 70% of the typical phenol-formaldehyde formulation could be replaced by the more environmentally friendly residues that are byproducts of ruminal-based fermentations. Even if it would not be possible to replace all of the phenol-formaldehyde adhesive, decreasing significant amounts of this material would provide for healthier composites by decreasing the amount of formaldehyde outgassing that are a human health concern[53]. Key to the effectiveness of fermentation residues is creating the correct balance of lignin, the blend of rumen microbes and the types

of glycocalyx material, and other minor phenolic materials in the plant materials.

This chapter is not meant to be a comprehensive assessment of biomass to biofuels, but rather a look at unconventional approaches that would enhance the sustainability of the entire process. To meet the goals of biofuel production by 2030 will require optimizing land use for food, feed, and bioenergy production. It should be approached from a standpoint of develop‐ ing a viable biofuel production system that increases the amount of energy stored in the molecules making up the biofuels, i.e., longer-chain molecules, more energy per unit of fuel. To be sustainable into the future we must be willing to develop alternative systems that supply a range of biomaterials. Although the producing energy alternatives is of major concern at the present time we should be evaluating and developing bioenergy systems that allow flexibility not only in terms of feedstock going in, but the products coming out. Development of biomass

**4. Conclusion**

562 Biofuels - Status and Perspective

Ronald Hatfield\*

Address all correspondence to<EMAIL_ADDRESS>
USDA-Agricultural Research Service, U.S. Dairy Forage Research Center, Madison, WI, USA