**The Application of Nanodiamond in Biotechnology and Tissue Engineering**

Lucie Bacakova, Antonin Broz, Jana Liskova, Lubica Stankova, Stepan Potocky and Alexander Kromka

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/63549

#### **Abstract**

Diamond in the allotrope form consists of carbon atoms arranged in a cubic crystal structure covalently bonded in sp3 hybridization. Diamond has emerged as a very promising material for various biomedical applications due to its excellent mechani‐ cal properties (hardness, low friction coefficient, good adhesiveness to the underlying substrate, good interlayer cohesion), optical properties (the ability to emit intrinsic luminescence), electrical properties (good insulator in the pristine state and semicon‐ ductor after doping), chemical resistance (low chemical reactivity and resistance to wet etching) and biocompatibility (little if any toxicity and immunogenicity). For ad‐ vanced biomedical applications, diamond is promising particularly in its nanostruc‐ tured forms, namely nanoparticles, nanostructured diamond films and composite scaffolds in which diamond nanoparticles are dispersed in a matrix (mainly nanodia‐ mond-loaded nanofibrous scaffolds). This chapter summarizes both our long-term experience and that of other research groups in studies focusing on the interaction of cells (particularly bone-derived cells) with nanodiamonds as nanoparticles, thin films and composites with synthetic polymers. Their potential applications in bioimaging, biosensing, drug delivery, biomaterial coating and tissue engineering are also reviewed.

**Keywords:** diamond nanoparticles, nanocrystalline diamond films, nanodiamondpolymer composites, bioimaging, biosensing, drug delivery, nanodiamond cytotoxici‐ ty, biomaterial coating, tissue engineering

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

#### **1. Introduction**

The approaching nano- and biotechnological era expects the implementation of "smart" and "functional" materials, which will be used not only as a mechanically passive substrates, but also as active devices (i.e., electronic or optical devices and sensors, micro-electro-mechanical systems, smart drug delivery systems, bioactive substrates for the cell adhesion and growth, etc.).

Diamond represents a class of perspective multifunctional materials with morphological, chemical, optical and electronic properties tailorable on demand for specific needs, especially for life science, tissue engineering or regenerative medicine.

The bonds between the carbon atoms in the diamond lattice are covalent and predominantly of sp3 hybridization [1, 2]. The diamond lattice of covalently bonded carbon atoms endows the diamond with extraordinary combination of intrinsic properties, particularly high hardness and thermal conductivity [3, 4]. The fracture toughness of the diamond has been measured to be 2 MPa m1/2, which is a relatively high value compared to other gemstones or other optical materials [5]. Diamond toughness is also dependent on the crystallographic plane on which the fracture force is expressed. The lowest cleavage energies were measured for the <111> plane [6]. The combination of the highest Young's modulus, a high fracture toughness, a low friction coefficient and a low thermal expansion predetermines it as a material for protective layers and coatings. From optical point of view, diamond offers a unique combination of transparency in most of the ultraviolet, visible and infrared regions (from 227 nm to far IR). The refractive index of diamond is 2.41–2.44 (656–486 nm). As a semiconductor, diamond has an excellent combination of properties (except of electron mobility) [7, 8]. Other remarkable properties of diamond include a high wear resistance [9–11], pressure resistance [12], strong adhesion to the underlying substrate, i.e., when deposited in the form of films [10, 13, 14], a low friction coefficient [15] and high chemical stability (e.g., low reactivity and resistance to wet etching [2]). These properties can be further improved by various techniques of material engineering, e.g., by modulating the synthesis and structure of diamond and by various combinations of diamond with other elements and compounds. For example, hardness, fracture toughness, Young's modulus and wear resistance of diamond can be markedly enhanced by the prepa‐ ration of diamond in the form of aggregated diamond nanorods [9], its hardness and thermal stability can be increased by nanotwinning with cubic boron nitride [4] and the adhesion of diamond films can be improved by laser treatment of the underlying substrate [14].

The optical properties of the diamond are also excellent. Pure diamond transmits visible light and appears as a clear colorless crystal. However, diamond is capable of high optical dispersion (i.e., the ability to disperse light of different colors). In addition, although the diamond lattice is very strong and rigid, it can contain defects or be contaminated with foreign atoms (N, B, H, Ni, Co, Cr, Si), e.g., during the growth of diamond lattice. These elements can also be introduced into this lattice by ion implantation. The lattice defects and the presence of foreign atoms are responsible for the various colors of diamond, e.g., yellow and orange (nitrogen), brown (nitrogen and lattice defects), blue (boron), green (trace amounts of nickel or radiation exposure), gray (the presence of boron or unsaturated forms of carbon), black (inclusions of graphite and iron) or also purple, pink and red (due to changes in the electron structure by plastic deformation during traveling of diamonds from the Earth's mantle to its surface *via* magma) [16]. The defects and the presence of foreign atoms in the diamond lattice also produce the intrinsic luminescence (fluorescence) of diamond [2, 17–19]. The electrical properties of diamond can be tuned by dopants. Pure diamonds can act as excellent electrical insulators but after chemical doping, e.g., the incorporation of boron, they are converted into a p-type semiconductor [20] or superconductor [21]. Boron is by far the most widely used dopant but other dopants are also possible [22].

**1. Introduction**

60 Diamond and Carbon Composites and Nanocomposites

etc.).

of sp3

The approaching nano- and biotechnological era expects the implementation of "smart" and "functional" materials, which will be used not only as a mechanically passive substrates, but also as active devices (i.e., electronic or optical devices and sensors, micro-electro-mechanical systems, smart drug delivery systems, bioactive substrates for the cell adhesion and growth,

Diamond represents a class of perspective multifunctional materials with morphological, chemical, optical and electronic properties tailorable on demand for specific needs, especially

The bonds between the carbon atoms in the diamond lattice are covalent and predominantly

diamond films can be improved by laser treatment of the underlying substrate [14].

The optical properties of the diamond are also excellent. Pure diamond transmits visible light and appears as a clear colorless crystal. However, diamond is capable of high optical dispersion (i.e., the ability to disperse light of different colors). In addition, although the diamond lattice is very strong and rigid, it can contain defects or be contaminated with foreign atoms (N, B, H, Ni, Co, Cr, Si), e.g., during the growth of diamond lattice. These elements can also be introduced into this lattice by ion implantation. The lattice defects and the presence of foreign atoms are responsible for the various colors of diamond, e.g., yellow and orange (nitrogen), brown (nitrogen and lattice defects), blue (boron), green (trace amounts of nickel or radiation exposure), gray (the presence of boron or unsaturated forms of carbon), black (inclusions of

 hybridization [1, 2]. The diamond lattice of covalently bonded carbon atoms endows the diamond with extraordinary combination of intrinsic properties, particularly high hardness and thermal conductivity [3, 4]. The fracture toughness of the diamond has been measured to be 2 MPa m1/2, which is a relatively high value compared to other gemstones or other optical materials [5]. Diamond toughness is also dependent on the crystallographic plane on which the fracture force is expressed. The lowest cleavage energies were measured for the <111> plane [6]. The combination of the highest Young's modulus, a high fracture toughness, a low friction coefficient and a low thermal expansion predetermines it as a material for protective layers and coatings. From optical point of view, diamond offers a unique combination of transparency in most of the ultraviolet, visible and infrared regions (from 227 nm to far IR). The refractive index of diamond is 2.41–2.44 (656–486 nm). As a semiconductor, diamond has an excellent combination of properties (except of electron mobility) [7, 8]. Other remarkable properties of diamond include a high wear resistance [9–11], pressure resistance [12], strong adhesion to the underlying substrate, i.e., when deposited in the form of films [10, 13, 14], a low friction coefficient [15] and high chemical stability (e.g., low reactivity and resistance to wet etching [2]). These properties can be further improved by various techniques of material engineering, e.g., by modulating the synthesis and structure of diamond and by various combinations of diamond with other elements and compounds. For example, hardness, fracture toughness, Young's modulus and wear resistance of diamond can be markedly enhanced by the prepa‐ ration of diamond in the form of aggregated diamond nanorods [9], its hardness and thermal stability can be increased by nanotwinning with cubic boron nitride [4] and the adhesion of

for life science, tissue engineering or regenerative medicine.

For all these outstanding properties, diamond is an attractive material for technical and industrial application, e.g., for polishing, machining, cutting and drilling tools. These tools are also useful for biomedical applications, e.g., for polishing biomaterials with diamond paste or cutting them with a diamond saw [23, 24], for bone resection or craniotomy using a diamond threadwire saw [25, 26] or for dental drilling in stomatology [27]. Other important fields for diamond application are electronics, optics or the jewelry trade.

This chapter will concentrate on newly explored applications of diamond in biotechnologies and life sciences, such as bioimaging, biosensing, tissue engineering, controlled drug and gene delivery and for detecting and capturing various biomolecules and coating body implants. For some of these uses, diamond is attractive mainly in its nanosized or nanostructured form, namely free nanoparticles (1D), planar films from nanodiamonds (2D), and composites of diamond nanoparticles (DNPs) and other molecules, particularly polymers in the form of 3D scaffolds, as shown in **Figure 1**.

**Figure 1** Schematic view of three diamond material forms representing free nanoparticles (1D), planar nanodiamond films (2D) and composite 3D scaffolds.

Our earlier studies included in this chapter were focused mainly on nanocrystalline diamond (NCD) films as substrates for the adhesion, growth and osteogenic differentiation of human bone-derived cells in the form of commercially available cell lines, namely human osteoblastlike MG 63 cells [20, 28–33] or Saos-2 cells [34–41], primary osteoblasts [42] and bone marrow mesenchymal stem cells (MSCs) [34, 42, 43]. The cell behavior on NCD films was further modulated by the size of the surface irregularities, measured by the root mean square (RMS) roughness parameter [34, 35], by the shape of these irregularities (nanocones and nanorods [36, 37], by termination of the NCD films with oxygen or hydrogen [38–42] and by doping of these films with boron [20, 32, 33]. NCD films micropatterned with H-terminated and Oterminated domains were also used for construction of biosensors [44]. Some of our studies were also dedicated to the construction of polymeric nanofibrous scaffolds reinforced with DNPs as potential scaffolds for bone tissue engineering. All the mentioned studies are summarized in **Table 1**.



**Table 1.** Summarization of our earlier studies on cell performance on NCD films and polymeric nanofibrous scaffolds loaded with diamond nanoparticles.

#### **2. Diamond nanoparticles**

mesenchymal stem cells (MSCs) [34, 42, 43]. The cell behavior on NCD films was further modulated by the size of the surface irregularities, measured by the root mean square (RMS) roughness parameter [34, 35], by the shape of these irregularities (nanocones and nanorods [36, 37], by termination of the NCD films with oxygen or hydrogen [38–42] and by doping of these films with boron [20, 32, 33]. NCD films micropatterned with H-terminated and Oterminated domains were also used for construction of biosensors [44]. Some of our studies were also dedicated to the construction of polymeric nanofibrous scaffolds reinforced with DNPs as potential scaffolds for bone tissue engineering. All the mentioned studies are

**References Topic Cell model**

MG 63

MG 63

MG 63

Saos-2 MSC

Saos-2

Saos-2

Saos-2

Saos-2

Saos-2

Saos-2

MG 63, endothelial CPAE cells

Cell adhesion and proliferation on nanostructured

Cell adhesion and proliferation on nanostructured

and hierarchically organized submicron- and nanostructured

and hierarchically organized submicron- and nanostructured

Adhesion, osteogenic cell differentiation and immune activation of cells on nanostructured and hierarchically organized submicron- and

Adhesion, proliferation, viability, mitochondrial activity and osteogenic differentiation of cells on nanostructured and hierarchically organized submicron- and nanostructured NCD

Regulation of the cell adhesion by the surface roughness

Regulation of the cell adhesion by the shape of surface

Adhesion and growth of cells on NCD films patterned with

Adsorption of fetal bovine serum proteins and cell adhesion on NCD films patterned with O-terminated and H-terminated

Adsorption of fibronectin and cell adhesion on NCD films patterned

O-terminated and H-terminated microdomains

with O-terminated and H-terminated microdomains

Regulation of the osteogenic cell differentiation by the surface

Regulation of the cell adhesion, activity of focal adhesion kinase and osteogenic cell differentiation by the shape of surface irregularities of the

summarized in **Table 1**.

NCD films

62 Diamond and Carbon Composites and Nanocomposites

NCD films

films

nanostructured NCD films

roughness of NCD films

irregularities of the NCD films

of NCD films

NCD films

microdomains

Bacakova *et al*. [28]

Grausova *et al*. [29]

Grausova *et al*. [30]

Grausova *et al*. [31]

Broz *et al*. [34]

Kalbacova *et al*. [35]

Babchenko *et al*. [37]

Kalbacova *et al*. [36]

Rezek *et al*. [38]

Rezek *et al*. [39]

Ukraintsev *et al*. [40] Individual DNPs can be advantageously used for drug and gene delivery, bioimaging technologies and biosensor construction. For these purposes, it is necessary to attach various atoms or molecules on the DNP surface, to achieve good dispersiveness of the DNPs in aqueous solutions, such as drug vehicles, buffers, physiological solution, cell culture media or body fluids and also to add chemical functionality to the DNPs in order to enable the uptake of DNPs by cells. DNPs have a high surface/volume ratio, i.e., they have a relatively large surface, on which not only atoms and small molecules, but also macromolecules (e.g., various drugs, nucleic acids and proteins) can be attached by physi- or chemisorption (for a review, see [47]). However, this attachment can be limited, for example, due to DNP hydrophobicity and tendency to aggregate in aqueous solutions. Thus, it is necessary to engineer the surface of DNPs for their specific applications [2]. The attachment of synthetic and biological molecules to the nanodiamond surface and the dispersion of DNPs in the aqueous media can be improved by various types of functionalization of the nanodiamond surface, e.g., by oxygen-containing chemical functional groups, such as –OH, −COOH [48–50], amine groups [48, 51] and by various biomolecules, such as lysine [52], biotin [53], thionine, trimethylamine [49], polygly‐

cerol [54, 55] and RGD-containing oligopeptides, i.e., ligands for adhesion receptors on the cell surface [54, 56]. This functionalization can also improve the penetration of DNPs into cells (i.e., by crossing cytoplasmic and nuclear membranes) or by their internalization through clathrinmediated endocytosis [57–59]. Functionalization of the DNPs with thiol groups (−SH) enables their conjugation with gold and silver nanoparticles [60, 61]. Such complexes are promising for photothermal therapy against tumors [60].

#### **2.1. Controlled drug and gene delivery**

Nanodiamond-based drug delivery has been developed particularly for advanced tumor therapies, e.g., the treatment of multidrug-chemoresistant leukemia. In experiments *in vitro*, daunorubicin conjugated with DNPs was efficient against a multidrug-resistant K562 human myelogenous leukemia cell line, which was able to overcome treatment with daunorubicin alone [62]. Other anticancer drugs, which were successfully conjugated with DNPs and delivered to cells, were 10-hydroxycamptothecin [63], polymyxin [64] or doxorubicin, an apoptosis-inducing drug widely used in chemotherapy [54, 65]; for a review, see [1, 66]. A nanodiamond-mediated doxorubicin delivery system was designed to inhibit the lung metastasis of breast cancer [67] and also to treat malignant brain gliomas [68]. The nanodia‐ mond-doxorubicin complexes were more significantly efficient in killing glioma cells *in vitro* and also *in vivo* in rats than the uncomplexed drug [68].

Other therapeutic molecules which can be delivered into cells through DNPs include cell growth and differentiation factors, antibodies, vaccines and anti-inflammatory drugs. For example, basic fibroblast growth factor (bFGF) and bone morphogenetic protein-2 (BMP-2) were loaded onto DNPs by physisorption, forming a stable colloidal solution. This solution enabled the release of both factors in slightly acidic conditions, induced the proliferation and osteogenic differentiation of osteoblast progenitor cells and was injectable. Thus, this solution was an effective alternative to the currently used delivery of bFGF and BMP-2 to the surgical site by the implantation of bulky collagen sponges [69]. Carboxylated DNPs conjugated with growth hormone were tested for a specific targeting growth hormone receptor in cancer cells and potential anticancer therapy [70].

DNPs can also serve as a stable delivery platform for therapeutic antibodies, as revealed by studies on the transforming growth factor β (TGF-β) antibody. The complexes of this antibody with DNPs were stable in water, but under physiological conditions simulated in serumsupplemented cell culture, the release of the active antibody was detected [71]. An interesting newly developed application of DNPs is delivering vaccines through the skin for painless and efficient immunization. These nanoparticles could improve limited drug delivery through the stratum corneum, which has a very dense structure and only allows the penetration of small molecules with a molecular weight of below 500 Da (for a review, see [72]). Fluorescent nanodiamonds functionalized by hyperbranched polyglycerol also proved to be promising carriers for delivering aluminum oxyhydroxide, i.e., a compound widely used as an immu‐ nologic adjuvant of vaccines [73].

For localized drug delivery, DNPs can be incorporated into various materials, e.g., films for potential implant coating [74]. For example, aqueous dispersible detonation nanodiamonds were assembled into a closely packed multilayer nanofilm with positively charged poly-Llysine *via* the layer-by-layer deposition technique, and integrated with drugs suppressing the release of inflammatory cytokines (tumor necrosis factor-alpha, interleukin-6) and nitric oxide synthase induced by lipopolysaccharides [65].

DNPs can also be effectively used for the proteolysis and digestion of glycopeptides. Trypsin and peptide-N-glycosidase F, immobilized on detonation nanodiamond, were more efficient than the free enzymes or commercially available beads immobilized with trypsin [75].

As for gene delivery, DNPs can act as excellent nonviral vectors for binding and transferring plasmid DNA and small interfering RNA (siRNA), particularly after functionalization, e.g., with –OH groups [49] or lysine [52]. DNPs, rendered cationic by coating with polyethylenei‐ mine or by termination with hydrogen in plasma, efficiently delivered siRNA into Ewing sarcoma cells and blocked the expression of EWS/FLI-1 oncogene in these cells. In addition, the association of EWS/FLI-1 silencing by the siRNA/nanodiamond complex with a vincristine treatment potentiated the cytotoxic effect of this chemotherapeutic drug [59]. Similarly, DNPs conjugated with RGD oligopeptides and siRNA for vascular endothelial growth factor (VEGF) can be used in antiangiogenic gene therapy to inhibit tumor growth *via* the downregulation of the VEGF expression [56]. Last but not least, an array of DNPs in the form of nanoneedles facilitated the delivery of double-strand DNAs (dsDNA90) into cells to activate the pathway involving the stimulator of interferon genes (STING). In addition, this array was successfully utilized to isolate the transcriptional factor, NF-kB, from intracellular regions without dam‐ aging the cells, upon STING activation [76].

#### **2.2. Bioimaging**

cerol [54, 55] and RGD-containing oligopeptides, i.e., ligands for adhesion receptors on the cell surface [54, 56]. This functionalization can also improve the penetration of DNPs into cells (i.e., by crossing cytoplasmic and nuclear membranes) or by their internalization through clathrinmediated endocytosis [57–59]. Functionalization of the DNPs with thiol groups (−SH) enables their conjugation with gold and silver nanoparticles [60, 61]. Such complexes are promising

Nanodiamond-based drug delivery has been developed particularly for advanced tumor therapies, e.g., the treatment of multidrug-chemoresistant leukemia. In experiments *in vitro*, daunorubicin conjugated with DNPs was efficient against a multidrug-resistant K562 human myelogenous leukemia cell line, which was able to overcome treatment with daunorubicin alone [62]. Other anticancer drugs, which were successfully conjugated with DNPs and delivered to cells, were 10-hydroxycamptothecin [63], polymyxin [64] or doxorubicin, an apoptosis-inducing drug widely used in chemotherapy [54, 65]; for a review, see [1, 66]. A nanodiamond-mediated doxorubicin delivery system was designed to inhibit the lung metastasis of breast cancer [67] and also to treat malignant brain gliomas [68]. The nanodia‐ mond-doxorubicin complexes were more significantly efficient in killing glioma cells *in vitro*

Other therapeutic molecules which can be delivered into cells through DNPs include cell growth and differentiation factors, antibodies, vaccines and anti-inflammatory drugs. For example, basic fibroblast growth factor (bFGF) and bone morphogenetic protein-2 (BMP-2) were loaded onto DNPs by physisorption, forming a stable colloidal solution. This solution enabled the release of both factors in slightly acidic conditions, induced the proliferation and osteogenic differentiation of osteoblast progenitor cells and was injectable. Thus, this solution was an effective alternative to the currently used delivery of bFGF and BMP-2 to the surgical site by the implantation of bulky collagen sponges [69]. Carboxylated DNPs conjugated with growth hormone were tested for a specific targeting growth hormone receptor in cancer cells

DNPs can also serve as a stable delivery platform for therapeutic antibodies, as revealed by studies on the transforming growth factor β (TGF-β) antibody. The complexes of this antibody with DNPs were stable in water, but under physiological conditions simulated in serumsupplemented cell culture, the release of the active antibody was detected [71]. An interesting newly developed application of DNPs is delivering vaccines through the skin for painless and efficient immunization. These nanoparticles could improve limited drug delivery through the stratum corneum, which has a very dense structure and only allows the penetration of small molecules with a molecular weight of below 500 Da (for a review, see [72]). Fluorescent nanodiamonds functionalized by hyperbranched polyglycerol also proved to be promising carriers for delivering aluminum oxyhydroxide, i.e., a compound widely used as an immu‐

For localized drug delivery, DNPs can be incorporated into various materials, e.g., films for potential implant coating [74]. For example, aqueous dispersible detonation nanodiamonds

for photothermal therapy against tumors [60].

and also *in vivo* in rats than the uncomplexed drug [68].

**2.1. Controlled drug and gene delivery**

64 Diamond and Carbon Composites and Nanocomposites

and potential anticancer therapy [70].

nologic adjuvant of vaccines [73].

Due to stable and controllable photoluminescence, DNPs are also promising for advanced bioimaging techniques. Like the various colors of diamonds, mentioned above, this photolu‐ minescence is based on crystallographic defects and particularly on optical centers created by the incorporation of foreign atoms (N, Si or Cr) into the diamond lattice. DNPs with a nitrogenvacancy (NV) color center were detectable in living cells *in vitro*, and also after application into chicken embryos *in vivo*, using a standard confocal microscope [19]. In addition, DNPs with NV color centers offer great potential for advanced imaging techniques, such as optical super resolution nanoscopy, magneto-optical (spin-assisted) subwavelength localization and imaging, single molecule spin imaging or nanoscale imaging of biomagnetic fields [17]. Photoluminescent nanodiamonds of size less than 50 nm were used to investigate the mech‐ anism of their uptake by living cells. By selectively blocking different uptake processes, it was shown that DNPs enter cells mainly by clathrin-mediated endocytosis. Intracellularly, the largest nanoparticles and aggregates were localized in vesicles, while the smallest particles appeared free in the cytosol [58]. Fluorescent DNPs were also used for tracking the trans‐ planted lung stem cells in mice with naphthalene-induced lung injury, and these stem cells were found at terminal bronchioles of the lungs for 7 days after intravenous transplantation [77].

The optical activity of the luminescent color centers in DNPs depends on their proximity to the nanoparticle surface and its surface termination [2]. The photoluminescence can also be modulated by changes in the environment surrounding the DNPs. For example, coating the DNPs with a polymer film resulted in a reduction of the lifetime of NV centers and an average enhancement in their emission rate [78]. For photoacoustic imaging in biological tissues, nanodiamonds with very high absorption in the near-infrared range can be created by irradiation with the He+ ion beam [79]. The fluorescence of DNPs can also be induced by surface passivation of the DNPs with bis(3-aminopropyl) terminated poly(ethylene glycol) [60]. A nanodiamond-polyglycerol-gadolinium(II) conjugate was designed and prepared as a novel nanodiamond-based magnetic resonance contrast agent dispersible in physiological media [55]. For combined bioimaging and drug delivery, photoluminescent DNPs were coated with a porous SiO2 shell [80].

#### **2.3. Biosensing**

DNPs with an NV color center are also emerging tools for nanoscale sensing, e.g., sensing molecular fluctuations and temperatures in live cellular environments, detecting and meas‐ uring magnetic and electric fields, and measuring ion concentrations and cell membrane potentials [17, 18]. DNPs immobilized onto a gold electrode by direct adsorption were used as a biosensor for determining the concentration of glucose and lactate [81]. Arrays of DNPs in the form of nanoneedles were applied for intracellular sensing, e.g., of intracellular signaling through NF-kB, translocated from the cell cytoplasm to the nucleus region [76]. Other sensoric applications of nanodiamonds require the creation of continuous diamond films, and they are discussed in Section 3 (Nanostructured diamond films).

#### **2.4. The potential cytotoxicity of diamond nanoparticles**

For all the applications mentioned above, it is necessary to use nontoxic nanoparticles. DNPs are generally considered as materials with little, if any cytotoxicity. These nanoparticles dispersed in the cell culture media did not affect the adhesion, growth, viability and differen‐ tiation of various cell types (for a review, see [82, 83]). For example, the addition of DNPs in the culture medium did not affect the adhesion, growth and adipogenic or osteogenic differ‐ entiation of human adipose tissue-derived stem cells [84]. Also the labeling of lung stem cells with fluorescent DNPs *in vivo* in mice did not affect the ability of these cells to grow and to differentiate into type I and type II pneumocytes [77].

In spite of these encouraging findings, the number of reports on nanodiamond cytotoxicity is increasing. The first report demonstrated the damage and destruction of human erythrocytes and leucocytes by DNPs [85, 86]. For low and moderate concentrations, DNPs did not negatively influence the viability of isolated inflammatory neutrophils and even enhanced their response to bacterial agents. However, the viability and antibacterial reaction of these cells was suppressed at high concentrations of DNPs (≥1 g/l; [87]). Similar reactions of neutrophils were also observed *in vivo* after intravenous and intraperitoneal administration of DNPs into rats [88]. DNPs evoked significant activation of human platelets, and when administered intravenously in mice, they induced widespread pulmonary thromboembolism [89]. DNPs had antiangiogenic effects, exerted by the downregulation of the gene and protein expression of the proangiogenic bFGF [90]. DNPs also induced apoptosis and necrosis of vascular endothelial cells [91, 92] and caused DNA damage of mouse embryonic stem cells [93] and human glioblastoma U87 cells [94]. Moreover, DNPs decreased the proliferation and viability of various cell types, e.g., human cervical carcinoma HeLa cells [95], human osteo‐ blast-like MG 63 cells and primary rat MSCs [96].

modulated by changes in the environment surrounding the DNPs. For example, coating the DNPs with a polymer film resulted in a reduction of the lifetime of NV centers and an average enhancement in their emission rate [78]. For photoacoustic imaging in biological tissues, nanodiamonds with very high absorption in the near-infrared range can be created by

passivation of the DNPs with bis(3-aminopropyl) terminated poly(ethylene glycol) [60]. A nanodiamond-polyglycerol-gadolinium(II) conjugate was designed and prepared as a novel nanodiamond-based magnetic resonance contrast agent dispersible in physiological media [55]. For combined bioimaging and drug delivery, photoluminescent DNPs were coated with

DNPs with an NV color center are also emerging tools for nanoscale sensing, e.g., sensing molecular fluctuations and temperatures in live cellular environments, detecting and meas‐ uring magnetic and electric fields, and measuring ion concentrations and cell membrane potentials [17, 18]. DNPs immobilized onto a gold electrode by direct adsorption were used as a biosensor for determining the concentration of glucose and lactate [81]. Arrays of DNPs in the form of nanoneedles were applied for intracellular sensing, e.g., of intracellular signaling through NF-kB, translocated from the cell cytoplasm to the nucleus region [76]. Other sensoric applications of nanodiamonds require the creation of continuous diamond films, and they are

For all the applications mentioned above, it is necessary to use nontoxic nanoparticles. DNPs are generally considered as materials with little, if any cytotoxicity. These nanoparticles dispersed in the cell culture media did not affect the adhesion, growth, viability and differen‐ tiation of various cell types (for a review, see [82, 83]). For example, the addition of DNPs in the culture medium did not affect the adhesion, growth and adipogenic or osteogenic differ‐ entiation of human adipose tissue-derived stem cells [84]. Also the labeling of lung stem cells with fluorescent DNPs *in vivo* in mice did not affect the ability of these cells to grow and to

In spite of these encouraging findings, the number of reports on nanodiamond cytotoxicity is increasing. The first report demonstrated the damage and destruction of human erythrocytes and leucocytes by DNPs [85, 86]. For low and moderate concentrations, DNPs did not negatively influence the viability of isolated inflammatory neutrophils and even enhanced their response to bacterial agents. However, the viability and antibacterial reaction of these cells was suppressed at high concentrations of DNPs (≥1 g/l; [87]). Similar reactions of neutrophils were also observed *in vivo* after intravenous and intraperitoneal administration of DNPs into rats [88]. DNPs evoked significant activation of human platelets, and when administered intravenously in mice, they induced widespread pulmonary thromboembolism [89]. DNPs had antiangiogenic effects, exerted by the downregulation of the gene and protein

discussed in Section 3 (Nanostructured diamond films).

**2.4. The potential cytotoxicity of diamond nanoparticles**

differentiate into type I and type II pneumocytes [77].

ion beam [79]. The fluorescence of DNPs can also be induced by surface

irradiation with the He+

66 Diamond and Carbon Composites and Nanocomposites

a porous SiO2 shell [80].

**2.3. Biosensing**

The adverse effects of DNPs have usually been attributed to the oxidative stress generated by these particles [92, 95, 97, 98]. DNPs created by detonation synthesis affected the cellular content of glutathione and the activities of the main antioxidant enzymes (superoxide dismu‐ tase, catalase, glutathione peroxidase, glutathione reductase and glutathione S-transferase) in human umbilical vein endothelial cells *in vitro* [98]. Furthermore, after intravenous and intraperitoneal administration into rats *in vivo*, DNPs increased the activity of superoxide dismutase and decreased the activity of glutathione reductase and glutathione peroxidase within erythrocytes [88]. In addition, DNPs increased the adsorption and delivery of a large amount of sodium ions into cells, which induced osmotic stresses, cell swelling, increase in the intracellular levels of calcium and severe cellular damage [97].

The cytotoxicity of DNPs depends on their origin, size, surface functionalization, tendency to form aggregates and the presence of impurities. In studies reporting the adverse effects of DNPs, these nanoparticles were often prepared by detonation [85, 86, 90, 93, 96, 97], while in studies indicating the nontoxic effects of DNPs on cell behavior, the DNPs were prepared by non-detonation methods, e.g., milling diamond microcrystals [73, 99]. The detonation DNPs, also called ultrananocrystalline detonation diamond (UDD), are generally of smaller size (in nanometers) than non-detonation DNPs (tens of nanometers), and this small size has been considered as the main determinant of nanoparticle cytotoxicity [91]. The detonation diamonds often contain impurities, such as other carbon allotropes [53], iridium [100], carbon soot and oxides and carbides, including those of iron, chromium, silicon, calcium, copper, potassium, titanium and sulfur (for a review, see [1]). In a study by Keremidarska et al. [96], the increased cytotoxicity of DNPs was associated with their smaller size and presence of impurities, particularly other carbon allotropes. Thus, for biomedical studies, detonation nanodiamonds need to be extensively purified. For the removal of nondiamond carbon, the following chemicals can be used: ozone-enriched air, liquid oxidants such as HNO3, HClO4 or different acid mixture under pressure; metal impurities can be removed by treatment with HCl, H2SO4 and its mixtures with HNO3 or potassium dichromate (for a review, see [1, 96]).

The functionalization influence of detonation DNPs on their biocompatibility is controversial. Detonation nanodiamonds covered by cobalt ions showed a protective effect against contact dermatitis in mini-pigs, induced by local exposure of the animal skin to these ions [101]. Detonation nanodiamonds with hyperbranched polyglycerol coating showed good cytocom‐ patibility with A549 cancer cells and U937 macrophages [54]. Detonation nanodiamonds functionalized with thiol groups and conjugated with gold nanoparticles did not show significant cytotoxicity to a human lung carcinoma cell line [61]. On the other hand, detonation DNPs terminated with oxygen showed higher cytotoxicity to mouse embryonic stem cells than pristine unmodified DNPs, although this cytotoxicity was lower than in multiwalled carbon nanotubes [93].

In our studies, we compared the potential cytotoxicity of DNPs from three different origins. DNPs fabricated in the Nano Carbon Research Institute (Japan) under the product name NanoAmando were prepared by a detonation method and were hydrogen-terminated. A part of these DNPs was oxygen-terminated by annealing at 450°C in air. The size of both unmodi‐ fied and O-terminated detonation DNPs was only about 5 nm. The DNPs, which were acquired from Microdiamant AG (Switzerland) named MSY (Monocrystalline SYnthetic), were pre‐ pared using a high pressure high temperature (HPHT) method and were oxygen-terminated due to postproduction acid treatment by the producer. These DNPs were not further modified and were provided in several sizes ranging from 18 to 210 nm. The DNPs acquired from Adámas Nanotechnologies, Inc. (Raleigh, U.S.A.) were prepared using the HPHT method and were already provided in a water colloid with a concentration of 1000 mg/l. The size of these nanoparticles was 40 nm. The Adámas DNPs had intrinsic red fluorescence based on NV defects in their diamond lattice, which was further activated by annealing in a vacuum at 800°C.

**Figure 2.** The number and mitochondrial activity (measured by MTS test) of human osteoblast-like Saos-2 cells in 3 day-old cultures exposed to NanoAmando (NA), MSY and Adámas (A) diamond nanoparticles suspended in cell cul‐ ture medium (concentrations from 10 to 1000 μg/ml) for 48 hours of cultivation. Control cells were cultivated in pure medium without diamond nanoparticles. Mean ± S.D. (standard deviation).

**Figure 3.** Micrographs of cells cultivated for 48 hours with Adámas DNPs Olympus IX 71, obj. 40×. Phase contrast im‐ age (A), red fluorescence of Adámas DNPs in grayscale (B), and merged images A and B (C).

The DNPs were dispersed by ultrasonication in water to obtain a colloid with a DNP concen‐ tration of 10,000 μg/ml, and added to 24-hour-old cultures of human osteoblast-like Saos-2 cells in concentrations of 10, 100 or 1000 μg/ml of the culture medium. After 48 hours, the cell number was evaluated by cell counting on recorded microphotographs using Image J software and correlated with the activity of mitochondrial enzymes measured by a commercially available MTS test (Promega). The lowest cell numbers and mitochondrial activity were obtained in cultures with unmodified NanoAmando DNPs, and these values decreased with increasing nanoparticle concentration (**Figure 2**). The cell numbers slightly improved in cultures with annealed NanoAmando DNPs, but they did not reach the average values obtained in MSY and Adámas DNPs. Only the mitochondrial activity of cells cultured with annealed NanoAmando DNPs reached similar values to the cells exposed to MSY and Adámas DNPs. The annealing of DNPs at a high temperature probably had a similar effect to the chemical purification of DNPs in oxidizing agents. The fluorescence of Adámas DNPs was detected inside the fixed cells after 2 days of cultivation using the standard epifluorescence microscope Olympus IX 71 (**Figure 3**).

#### **3. Nanostructured diamond films**

In our studies, we compared the potential cytotoxicity of DNPs from three different origins. DNPs fabricated in the Nano Carbon Research Institute (Japan) under the product name NanoAmando were prepared by a detonation method and were hydrogen-terminated. A part of these DNPs was oxygen-terminated by annealing at 450°C in air. The size of both unmodi‐ fied and O-terminated detonation DNPs was only about 5 nm. The DNPs, which were acquired from Microdiamant AG (Switzerland) named MSY (Monocrystalline SYnthetic), were pre‐ pared using a high pressure high temperature (HPHT) method and were oxygen-terminated due to postproduction acid treatment by the producer. These DNPs were not further modified and were provided in several sizes ranging from 18 to 210 nm. The DNPs acquired from Adámas Nanotechnologies, Inc. (Raleigh, U.S.A.) were prepared using the HPHT method and were already provided in a water colloid with a concentration of 1000 mg/l. The size of these nanoparticles was 40 nm. The Adámas DNPs had intrinsic red fluorescence based on NV defects in their diamond lattice, which was further activated by annealing in a vacuum at

**Figure 2.** The number and mitochondrial activity (measured by MTS test) of human osteoblast-like Saos-2 cells in 3 day-old cultures exposed to NanoAmando (NA), MSY and Adámas (A) diamond nanoparticles suspended in cell cul‐ ture medium (concentrations from 10 to 1000 μg/ml) for 48 hours of cultivation. Control cells were cultivated in pure

**Figure 3.** Micrographs of cells cultivated for 48 hours with Adámas DNPs Olympus IX 71, obj. 40×. Phase contrast im‐

The DNPs were dispersed by ultrasonication in water to obtain a colloid with a DNP concen‐ tration of 10,000 μg/ml, and added to 24-hour-old cultures of human osteoblast-like Saos-2 cells in concentrations of 10, 100 or 1000 μg/ml of the culture medium. After 48 hours, the cell number was evaluated by cell counting on recorded microphotographs using Image J software

age (A), red fluorescence of Adámas DNPs in grayscale (B), and merged images A and B (C).

medium without diamond nanoparticles. Mean ± S.D. (standard deviation).

800°C.

68 Diamond and Carbon Composites and Nanocomposites

#### **3.1. Diamond films as substrates for cell adhesion and growth**

Our earlier studies and studies by other authors showed that NCD films provided excellent substrates for the adhesion, growth and phenotypic maturation of various cell types, such as neuronal and glial cells [102–104], epithelial cells [105], vascular endothelial cells [29], fibro‐ blasts [106] and particularly bone-derived cells (for a review, see [82, 83, 107]). The latter cells included commercially available lines of osteoblast-like cells (MG 63 [20, 28–33], Saos-2 [30, 31, 35, 41, 108]), primary human osteoblasts [42] and human bone marrow MSCs [42, 108, 109].

The positive effects of NCD films on cell performance can be explained by their nanostructure, i.e., the presence of irregularities at the nanoscale level on their surface (e.g., an average roughness Ra parameter equal or less than 100 nm). On these surfaces, the cell adhesionmediating proteins present in biological fluids (e.g., fibronectin and vitronectin in the serum supplement of the culture media) are adsorbed in a favorable geometrical conformation enabling the exposure of specific adhesion sites in these proteins (e.g., RGD motifs) to the cell adhesion receptors, i.e., integrins. In addition, it is believed that the nanostructured surfaces preferentially adsorb vitronectin (due to its relatively small and linear molecule), which is favorably recognized by osteoblasts through its KRSR amino acid sequence [110, 111] (for a review, see [107, 112]). Thus, the NCD films are promising particularly for the surface coating of bone and dental implants. In accordance with this concept, submicron crystalline diamond films (grain sizes 200–1000 nm) and particularly microcrystalline diamond films (grain size up to 2 μm) supported the adhesion and spreading of osteoblasts to a lesser extent than NCD films [113, 114]. Also in our earlier study, the number of human osteoblast-like Saos-2 cells adhering to NCD films deposited on silicone substrates of various roughness (RMS of 20, 270 and 500 nm) decreased with the increasing substrate roughness [34]. In addition, the NCD films of RMS of 20 nm and 270 nm supported better the osteogenic differentiation of Saos-2 cells than the surfaces of a RMS of 500 nm, as manifested by a higher activity of alkaline phosphatase and a higher cellular content of calcium and phosphates [35]. Similarly, on NCD films with hier‐ archically-organized submicron and nanoscale surface roughness (RMS of 301.0 nm and 7.6 nm, respectively), the MG 63 cells adhered in lower initial numbers than on nanorough only surfaces (RMS 8.2 nm), although the subsequent cell proliferation was higher on the hierarch‐ ically-organized surfaces [28].

Not only the size but also the shape of the surface irregularities on the diamond films is an important factor regulating the cell behavior. For example, in our earlier studies, two different types of nanoscale features, namely nanocones and nanorods, were prepared by plasma etching of nanodiamond films using Au and Ni masks, respectively. As revealed by immu‐ nofluorescence staining of vinculin, Saos-2 cells cultivated on relatively short and broad nanocones (height 5–100 nm, diameter up to 80 nm) formed less numerous but large focal adhesions, while the cells cultivated on relatively high and thin nanorods (height 120–200 nm, diameter 20–40 nm) formed more numerous, but very thin and fine focal adhesions [37]. The large focal adhesions on nanocones were associated with stronger cell adhesion, increased activation of focal adhesion kinase (FAK), and were better predestined for osteogenic cell differentiation [36].

In contrast to free DNPs, to the best of our knowledge, no study reported the cytotoxicity of NCD films. This may be due to the fact that NCD films are compact, chemically and mechan‐ ically stable and their surface is much better defined/controlled than the surface of DNPs. Moreover, planar NCD films do not interfere and/or penetrate into the cells. The surface termination of NanoAmando DNPs controlled by high-temperature treatment was shown as a crucial parameter for their cytotoxicity (see Section 3.4). NCD films were not cytotoxic even if NanoAmando DNPs were used for the substrate seeding prior to NCD deposition.

NCD films have usually been synthesized from methane or fullerene C60 precursors using the microwave plasma-enhanced chemical vapor deposition (MW PECVD) method. NCD films have been deposited on various substrates such as the silicon, glass or metals currently used in orthopedics and stomatology, e.g., Ti [115] and Ti-6Al-4V [116] (for a review, see [82, 83]). On these substrates, the NCD films behave as well-adhering, highly cohesive and mechanically and chemically resistant coatings. Not only were their potential adverse effects suppressed, but no particles were released from the compact films, which in addition behaved as a passivation barrier for particles and ions from the underlying materials. For example, silicon wafers, which are currently being used as experimental substrates for NCD deposition and behaved as cytotoxic for human osteoblast-like MG 63 cells and vascular endothelial CPAE cells, were rendered highly compatible with these cells by their continuous and hermetic encapsulation with NCD films [29].

Cell performance on NCD films can be further modulated by their functionalization with various atoms, chemical functional groups or biomolecules. For example, the adhesion, growth and osteogenic differentiation of bone-derived cells were better on O-terminated than on Hterminated NCD films. Human osteoblast-like Saos-2 cells, primary human osteoblasts and bone marrow MSCs in cultures on O-terminated NCD films showed better-developed focal adhesion plaques (**Figure 4**), reached higher cell population densities and produced more collagen I and alkaline phosphatase (i.e., an early and middle marker of osteogenic cell differentiation, respectively). The cells on O-terminated surfaces also showed a higher activity of alkaline phosphatase, i.e., an enzyme participating in bone matrix mineralization, and accordingly with it, these cells produced more calcium in their extracellular matrix (ECM) [41, 42]. Furthermore, human dental stem cells cultured on O-terminated diamond films deposited more ECM than on H-terminated films, and this matrix contained higher levels of calcium, oxygen and phosphorus [117]. Similar behavior has also been observed in human renal epithelial HK-2 cells cultured on borosilicate glass with NCD films. On the films terminated with O, these cells adhered and grew better than on H-terminated NCD films and uncoated borosilicate glass [105]. When NCD films were constructed as micropatterned films, i.e., with domains terminated with O or H, e.g., in the form of stripes of 30–200 μm in width, the cells adhered and grew preferentially on the O-terminated domains [38, 117].

nm, respectively), the MG 63 cells adhered in lower initial numbers than on nanorough only surfaces (RMS 8.2 nm), although the subsequent cell proliferation was higher on the hierarch‐

Not only the size but also the shape of the surface irregularities on the diamond films is an important factor regulating the cell behavior. For example, in our earlier studies, two different types of nanoscale features, namely nanocones and nanorods, were prepared by plasma etching of nanodiamond films using Au and Ni masks, respectively. As revealed by immu‐ nofluorescence staining of vinculin, Saos-2 cells cultivated on relatively short and broad nanocones (height 5–100 nm, diameter up to 80 nm) formed less numerous but large focal adhesions, while the cells cultivated on relatively high and thin nanorods (height 120–200 nm, diameter 20–40 nm) formed more numerous, but very thin and fine focal adhesions [37]. The large focal adhesions on nanocones were associated with stronger cell adhesion, increased activation of focal adhesion kinase (FAK), and were better predestined for osteogenic cell

In contrast to free DNPs, to the best of our knowledge, no study reported the cytotoxicity of NCD films. This may be due to the fact that NCD films are compact, chemically and mechan‐ ically stable and their surface is much better defined/controlled than the surface of DNPs. Moreover, planar NCD films do not interfere and/or penetrate into the cells. The surface termination of NanoAmando DNPs controlled by high-temperature treatment was shown as a crucial parameter for their cytotoxicity (see Section 3.4). NCD films were not cytotoxic even

NCD films have usually been synthesized from methane or fullerene C60 precursors using the microwave plasma-enhanced chemical vapor deposition (MW PECVD) method. NCD films have been deposited on various substrates such as the silicon, glass or metals currently used in orthopedics and stomatology, e.g., Ti [115] and Ti-6Al-4V [116] (for a review, see [82, 83]). On these substrates, the NCD films behave as well-adhering, highly cohesive and mechanically and chemically resistant coatings. Not only were their potential adverse effects suppressed, but no particles were released from the compact films, which in addition behaved as a passivation barrier for particles and ions from the underlying materials. For example, silicon wafers, which are currently being used as experimental substrates for NCD deposition and behaved as cytotoxic for human osteoblast-like MG 63 cells and vascular endothelial CPAE cells, were rendered highly compatible with these cells by their continuous and hermetic

Cell performance on NCD films can be further modulated by their functionalization with various atoms, chemical functional groups or biomolecules. For example, the adhesion, growth and osteogenic differentiation of bone-derived cells were better on O-terminated than on Hterminated NCD films. Human osteoblast-like Saos-2 cells, primary human osteoblasts and bone marrow MSCs in cultures on O-terminated NCD films showed better-developed focal adhesion plaques (**Figure 4**), reached higher cell population densities and produced more collagen I and alkaline phosphatase (i.e., an early and middle marker of osteogenic cell differentiation, respectively). The cells on O-terminated surfaces also showed a higher activity of alkaline phosphatase, i.e., an enzyme participating in bone matrix mineralization, and

if NanoAmando DNPs were used for the substrate seeding prior to NCD deposition.

ically-organized surfaces [28].

70 Diamond and Carbon Composites and Nanocomposites

differentiation [36].

encapsulation with NCD films [29].

**Figure 4.** The immunofluorescence of paxillin (green), a protein of focal adhesion plaques, in primary human osteo‐ blasts on day 7 after seeding on H-terminated (Diam. H) and O-terminated (Diam. O) NCD films. Actin cytoskeleton is stained with Phalloidin-Texas Red, cell nuclei with Hoechst #33342 (blue). Bar = 25 μm.

The positive influence of O-terminated diamond films on cell behavior has been explained by the higher polar component of the free surface energy and higher wettability of these films [118] (for a review, see [41]). It is known that wettable surfaces (similar to the nanostructured surfaces mentioned above) promote the adsorption of cell adhesion-mediating molecules in a bioactive and flexible geometrical conformation, and their effective recognition and binding by the cell adhesion receptors (for a review, see [112]). Our earlier study showed that the fetal bovine serum (FBS), i.e., an important supplement of standard cell culture media and an important source of cell adhesion-mediating proteins (e.g., vitronectin, fibronectin), formed more compact and better-adherent films on O-terminated NCD surfaces than on H-terminated NCD, where the FBS layer tended to peel from the material surface [40]. The following experiments with preadsorption of NCD films with O-terminated and H-terminated micro‐ domains with FBS proteins revealed that the preferential growth of Saos-2 cells on O-termi‐ nated domains is mediated by fibronectin [39].

Positive effects on surface polarity, wettability and cell adhesion were also observed in NCD films terminated with amine groups, which upregulated the adhesion of human femur osteoblasts [115, 118], or supported the adhesion of dorsal root ganglion neurons and Schwann cells isolated from Wistar rats and NG108-15 neuroblastoma-glioma hybrid cells [102].

However, in the case of the neural stem cells derived from mouse embryos, both the Oterminated and H-terminated NCD surfaces effectively supported well the cell proliferation, and this proliferation was even slightly better on H-terminated surfaces. Moreover, Hterminated NCD films were able to spontaneously induce (i.e., without the presence of differentiation factors in the culture medium) differentiation of the stem cells into neurons, while on O-terminated NCD films, the cells preferentially differentiated into oligodendrocytes [103]. The neuronal cell differentiation on H-terminated NCD films was explained by the high accumulation of fibronectin on these films, which then activated a signaling pathway involving β1-integrin adhesion receptors, FAK and mitogen-activated protein kinase/extracellular signaling-regulated kinase1/2 (MAPK/Erk1/2) [104]. Thus, the H-terminated surfaces seemed to have a greater potential to promote the regeneration of neurons in a damaged central nervous system (which is considered as impossible in human patients) than O-terminated surfaces.

Thus, the results of cell performance on H-terminated and O-terminated NCD films, obtained by various authors, are controversial. This disproportion can be attributed to different NCD qualities (the presence of non-diamond carbon phases, different grain size, etc.), the different reactivity of various cell types (neural, epithelial, osteogenic) or different cultivation condi‐ tions. For example, the studies by Chen et al. [103, 104] were performed using a medium with a low concentration of serum (only 2% of FBS), while the experiments with osteogenic cells were generally carried out in standard cell culture media supplemented with 10–15% of FBS. When osteogenic cells were cultured in a serum-free medium, no differences in cell adhesion on O- and H-terminated NCD domains were observed [38]. Another explanation could be the relatively high hydrophilicity of O-terminated surfaces (water drop contact angle ~20° *vs*. ~78° for H-terminated surfaces) in studies by Chen et al. [103, 104]. It is generally known that cell adhesion and growth is optimal on moderately hydrophilic surfaces (for a review, see [112]). On highly hydrophilic surfaces, the adsorption of cell adhesion-mediating proteins is weak and unstable, which hamper or disable cell adhesion. For example, extremely hydrophilic Oterminated nanostructured diamond surfaces (contact angle ~2°) almost completely resisted the adhesion of human bone marrow MSCs, whereas less hydrophilic H-terminated nanodia‐ mond surfaces (contact angle ~86°) gave good support for the adhesion and growth of these cells [109]. Also in our studies, human osteoblast-like MG 63 cells on O-terminated NCD films (contact angle about 20–35°) showed a higher tendency for spontaneous detachment at later culture intervals, when the surfaces had to hold a large number of cells [29–31] than the cells on H-terminated NCD films (contact angle about 85–90° [33], (for a review, see [112].

Another important modification of NCD films influencing cell behavior is the boron doping of these films. Lower concentrations of boron (approximately 100–1000 ppm of B) supported the proliferation and early osteogenic differentiation of human osteoblast-like MG 63 cells, which was manifested by an increased cell number and increased concentration of collagen I and alkaline phosphatase in the cells. Higher concentrations of boron (6700 ppm of B) enhanced cell adhesion and osteogenic cell differentiation, manifested by increased concentrations of vinculin, a focal adhesion protein, and osteocalcin, a calcium-binding ECM glycoprotein. These effects were probably due to the increased electrical conductivity of NCD films after boron doping [32]. It is known that electroactive materials are able to enhance the adhesion, growth and phenotypic maturation on various cell types, even without active stimulation with an electrical current (for a review, see [83, 112]).

Also, composite apatite-nanodiamond coatings were shown to improve the performance of bone cells. When electrodeposited on stainless steel, these coatings markedly enhanced the attachment, spreading and formation of focal adhesion plaques in osteoblast-like MG 63 cells, compared to the pure stainless steel and apatite coatings without nanodiamond, which was mediated by an increased adsorption of fibronectin on the composite coatings and its deposi‐ tion by the cells [119].

#### **3.2. Diamond films as substrates for biosensing and localized drug delivery**

osteoblasts [115, 118], or supported the adhesion of dorsal root ganglion neurons and Schwann cells isolated from Wistar rats and NG108-15 neuroblastoma-glioma hybrid cells [102].

However, in the case of the neural stem cells derived from mouse embryos, both the Oterminated and H-terminated NCD surfaces effectively supported well the cell proliferation, and this proliferation was even slightly better on H-terminated surfaces. Moreover, Hterminated NCD films were able to spontaneously induce (i.e., without the presence of differentiation factors in the culture medium) differentiation of the stem cells into neurons, while on O-terminated NCD films, the cells preferentially differentiated into oligodendrocytes [103]. The neuronal cell differentiation on H-terminated NCD films was explained by the high accumulation of fibronectin on these films, which then activated a signaling pathway involving β1-integrin adhesion receptors, FAK and mitogen-activated protein kinase/extracellular signaling-regulated kinase1/2 (MAPK/Erk1/2) [104]. Thus, the H-terminated surfaces seemed to have a greater potential to promote the regeneration of neurons in a damaged central nervous system (which is considered as impossible in human patients) than O-terminated

Thus, the results of cell performance on H-terminated and O-terminated NCD films, obtained by various authors, are controversial. This disproportion can be attributed to different NCD qualities (the presence of non-diamond carbon phases, different grain size, etc.), the different reactivity of various cell types (neural, epithelial, osteogenic) or different cultivation condi‐ tions. For example, the studies by Chen et al. [103, 104] were performed using a medium with a low concentration of serum (only 2% of FBS), while the experiments with osteogenic cells were generally carried out in standard cell culture media supplemented with 10–15% of FBS. When osteogenic cells were cultured in a serum-free medium, no differences in cell adhesion on O- and H-terminated NCD domains were observed [38]. Another explanation could be the relatively high hydrophilicity of O-terminated surfaces (water drop contact angle ~20° *vs*. ~78° for H-terminated surfaces) in studies by Chen et al. [103, 104]. It is generally known that cell adhesion and growth is optimal on moderately hydrophilic surfaces (for a review, see [112]). On highly hydrophilic surfaces, the adsorption of cell adhesion-mediating proteins is weak and unstable, which hamper or disable cell adhesion. For example, extremely hydrophilic Oterminated nanostructured diamond surfaces (contact angle ~2°) almost completely resisted the adhesion of human bone marrow MSCs, whereas less hydrophilic H-terminated nanodia‐ mond surfaces (contact angle ~86°) gave good support for the adhesion and growth of these cells [109]. Also in our studies, human osteoblast-like MG 63 cells on O-terminated NCD films (contact angle about 20–35°) showed a higher tendency for spontaneous detachment at later culture intervals, when the surfaces had to hold a large number of cells [29–31] than the cells

on H-terminated NCD films (contact angle about 85–90° [33], (for a review, see [112].

Another important modification of NCD films influencing cell behavior is the boron doping of these films. Lower concentrations of boron (approximately 100–1000 ppm of B) supported the proliferation and early osteogenic differentiation of human osteoblast-like MG 63 cells, which was manifested by an increased cell number and increased concentration of collagen I and alkaline phosphatase in the cells. Higher concentrations of boron (6700 ppm of B) enhanced cell adhesion and osteogenic cell differentiation, manifested by increased concentrations of

surfaces.

72 Diamond and Carbon Composites and Nanocomposites

Nanocrystalline and polycrystalline diamond films are also important for constructing biosensors. For example, NCD films patterned with O- and H-terminated microdomains were used for the construction of an impedance-based sensor for real-time monitoring of cell growth, and successfully applied in cultures of human osteoblast-like MG 63 cells [44]. An aptasensor for platelet-derived growth factor (PDGF) was designed on a NCD surface functionalized with carboxylic and amine groups, which served as units for immobilizing PDGF-binding aptamers [120]. Arrays based on polycrystalline diamond surfaces grafted with odorant binding proteins (i.e., small soluble proteins present in olfactory systems) were fabricated for potential artificial olfaction applications [121]. Electrodes coated with boron-doped diamond (BDD) films are also excellent materials for the construction of biosensors, especially after further modifications with other nanoparticles and molecules [122]. For example, a BDD electrode modified with silver nanoparticles was developed as a cholesterol sensor [123]. A biosensor for the detection of L-serine was fabricated using a BDD electrode modified with polycrystalline nickel and nickel(II) oxide (Ni-NiO) half nanotubes [124]. A BDD electrode modified with platinum nanoparticles (serving as a highly conductive catalytic transducer, and coupled with a competitive magneto-enzyme immunoassay), was used for the detection and degradation of the pesticide atrazine [125]. An electrochemical sensor for glutamate, an important neuro‐ transmitter in the mammalian central nervous system, was fabricated on doped chemical vapor deposition diamond electrodes and graphene nanoplatelet structures [126]. Diamond coatings also enhanced the sensitivity and operation range of optical fiber sensors [127].

As for the drug delivery, oxidized ultrananocrystalline diamond thin films were functionalized with type I collagen and an anti-inflammatory drug dexamethasone, in order to fabricate a hybrid therapeutically active substrate for localized drug delivery [128].

#### **4. Polymer-nanodiamond composites**

Another promising group of materials for biomedical applications, particularly for bone tissue engineering, is polymer-nanodiamond composites. Polymers in general are too soft and elastic for bone tissue engineering, and thus they require reinforcements with harder and stronger materials, e.g., ceramic, metallic or carbon nanoparticles, including DNPs. Polymers reinforced with DNPs have been used in the form of thin films, e.g., poly(L-lactide) (PLLA) films [129, 130], or as electrospun nanofibrous scaffolds made of poly(lactide-co-glycolide) (PLGA) [43, 45]. Both types of composites showed improved mechanical properties and effectively supported the adhesion and growth of osteoblast-like MG 63 cells and human bone marrow MSCs (**Figure 5**). In the case of PLLA films, fluorescent DNPs were used in order to localize these materials in tissues after implantation and to investigate their behavior *in vivo* [129, 130]. In the case of nanofibrous PLGA scaffolds, the DNP-loaded scaffolds supported the growth of MG 63 cells to a similar degree as the pure scaffolds, but accelerated the growth of MSCs [43, 45]. This might be due to the fact that the primary or low-passaged cells, such as MSCs, were more sensitive to the physical and chemical properties of their growth substrate than the cell lines, which were better adapted to the *in vitro* conditions.

**Figure 5.** The morphology of pure PLGA nanofibrous scaffolds (A), PLGA scaffolds loaded with ~23 wt.% of diamond nanoparticles (B) and human bone marrow mesenchymal stem cells on day 9 after seeding on the PLGA (C) or PLGAnanodiamond (D) scaffolds. A, B: scanning electron microscope, bar = 10 μm. C, D: cell membrane and cytoplasm stained with Texas Red C2-Maleimide; cell nuclei counterstained with Hoechst #33342. Leica SPE confocal microscope, bar = 20 μm.

However, nanofibrous PLLA-nanodiamond scaffolds exerted rather adverse effects on their colonization with bone cells. The number, mitochondrial activity and expression of some markers of osteogenic differentiation at the mRNA and protein level in the human osteoblastlike MG 63 and Saos-2 cells cultured on these scaffolds decreased with the increasing concen‐ tration of DNPs in these scaffolds (from about 0.4 to 12 wt.%; [46]). At the same time, the concentration of DNPs in nanofibrous PLGA-nanodiamond scaffolds was considerably higher, i.e., ~23 wt.% [43, 45]. The different cell behavior on the PLGA- and PLLA-based scaffolds was probably due to the different origin and properties of the DNPs used for loading these scaffolds. For PLGA scaffolds, the DNPs were prepared using a radio frequency plasma activated chemical vapor deposition (RF PACVD) method [131], while for the PLLA scaffolds, pristine NanoAmando DNPs (see Section 2.4) were used. These DNPs were much smaller, hydrophobic, and induced cytotoxic effects in Saos-2 cells when added into the cell culture medium. They were probably released from the scaffolds, and may have also increased the hydrophobicity of the scaffolds, which could negatively affect cell functions.

DNPs have been also added in materials other than polymers. For example, these nanoparticles were used to increase the corrosion resistance of Mg, i.e., a material promising to construct biodegradable bone implants. Specifically, the corrosion rate of Mg was reduced by 4.5 times, when 5 wt.% of DNPs were uniformly dispersed in the Mg matrix. At the same time, L-929 fibroblasts cultured on the Mg-nanodiamond composites maintained high cell viability and a healthy morphology [132].

#### **5. Conclusion**

for bone tissue engineering, and thus they require reinforcements with harder and stronger materials, e.g., ceramic, metallic or carbon nanoparticles, including DNPs. Polymers reinforced with DNPs have been used in the form of thin films, e.g., poly(L-lactide) (PLLA) films [129, 130], or as electrospun nanofibrous scaffolds made of poly(lactide-co-glycolide) (PLGA) [43, 45]. Both types of composites showed improved mechanical properties and effectively supported the adhesion and growth of osteoblast-like MG 63 cells and human bone marrow MSCs (**Figure 5**). In the case of PLLA films, fluorescent DNPs were used in order to localize these materials in tissues after implantation and to investigate their behavior *in vivo* [129, 130]. In the case of nanofibrous PLGA scaffolds, the DNP-loaded scaffolds supported the growth of MG 63 cells to a similar degree as the pure scaffolds, but accelerated the growth of MSCs [43, 45]. This might be due to the fact that the primary or low-passaged cells, such as MSCs, were more sensitive to the physical and chemical properties of their growth substrate

**Figure 5.** The morphology of pure PLGA nanofibrous scaffolds (A), PLGA scaffolds loaded with ~23 wt.% of diamond nanoparticles (B) and human bone marrow mesenchymal stem cells on day 9 after seeding on the PLGA (C) or PLGAnanodiamond (D) scaffolds. A, B: scanning electron microscope, bar = 10 μm. C, D: cell membrane and cytoplasm stained with Texas Red C2-Maleimide; cell nuclei counterstained with Hoechst #33342. Leica SPE confocal microscope,

However, nanofibrous PLLA-nanodiamond scaffolds exerted rather adverse effects on their colonization with bone cells. The number, mitochondrial activity and expression of some markers of osteogenic differentiation at the mRNA and protein level in the human osteoblastlike MG 63 and Saos-2 cells cultured on these scaffolds decreased with the increasing concen‐

bar = 20 μm.

than the cell lines, which were better adapted to the *in vitro* conditions.

74 Diamond and Carbon Composites and Nanocomposites

It can be concluded that diamond in the form of nanoparticles, nanostructured films and composite scaffolds, particularly nanofibrous scaffolds loaded with DNPs, is an exceptionally promising material for a wide range of biomedical applications. DNPs are suitable for drug and gene delivery, bioimaging, biosensing and also as additives to polymeric scaffolds for bone tissue engineering, particularly those nanofibrous, in order to improve their mechanical properties and increase their bioactivity. Nanodiamond films can be applied for localized drug delivery, construction of biosensors and particularly for bone implant coating due to their favorable effect on cell adhesion, growth and osteogenic differentiation. The cell performance on the nanodiamond films can be modulated by their roughness and topography, e.g., the size and shape of the surface irregularities, by termination with various atoms (e.g., O and H), chemical functional groups and biomolecules and by boron doping. However, free DNPs and DNPs dispersed in a material matrix should be applied with caution because of their higher reactivity with the surrounding environment and potential penetration into cells, which may finally result in their cytotoxicity.

#### **Acknowledgements**

Our studies included in this review were supported by the Agency for Healthcare Research, Ministry of Health of the Czech Republic (grant no. 15-32497A) and the Grant Agency of the Czech Republic (grant no. 14-04790S). Mrs. Paula Solon is gratefully acknowledged for her language revision of the manuscript.

#### **Author details**

Lucie Bacakova1\*, Antonin Broz1 , Jana Liskova1 , Lubica Stankova1 , Stepan Potocky2 and Alexander Kromka2


2 Institute of Physics of the Czech Academy of Sciences, v.v.i., Prague, Czech Republic

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**Author details**

Alexander Kromka2

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Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/63417

#### **Abstract**

The chapter presents the results of experimental investigation of the machinability of difficult to cut materials, e.g. silicon carbide particulate, aluminium metal matrix composite, cemented carbides layers, technical ceramics (Si3N4). The work focuses on the study of machinability during conventional and laser-assisted turning. Nowa‐ days, the difficult to cut materials have a great application in different fields of industry, e.g. automobile or aerospace. Metal matrix composites are known as the difficult to machine materials, because of the hardness and abrasive nature element—in this case —SiC.

The work reported concentrates on the machinability improvement of difficult to cut materials by laser-assisted machining (LAM) when compared with conventional cutting process. Influence of laser beam during laser-assisted turning on temperature of the heating area, tool wear, and machined surfaces' roughness was investigated. This research was carried out for polycrystalline diamond wedges (PCD), ceramics and sintered carbide wedges with coatings and uncoated ones. The results obtained with the laser assisted machining were compared with results obtained in conventional turning.

**Keywords:** laser-assisted machining, polycrystalline diamond, technical ceramics, metal matrix composites, cemented carbides

#### **1. Introduction**

The metal matrix composites as well as Si3N4 ceramics are important materials with many engineering and medical applications. However, the high hardness (2125HV) and brittleness of ceramics and the abrasive characteristics of the reinforced particulates in metal matrix composites (MMC) make serious difficulties when machining [1–3].

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

Mechanical and physical properties like high thermal resistance, high hardness, good corrosive resistance and chemical stability have encouraged use of ceramics and SiC reinforced alumi‐ nium-based composites in several engineering applications. The brittle and hard nature of these ceramics makes them difficult to machine using conventional techniques and damage caused to the surface while machining affects efficiency of components. Laser-assisted machining (LAM) has recently emerged as a potential technique for attaining high material removal rates.

MMCs and silicon nitride are increasingly attractive for engineering applications because of their superior combination of properties such as high strength, hardness, good corrosive resistance, and wear resistance. These materials are finding potential applications in the manufacture of automotive components such as some elements of cars, aerospace and medical [4, 5]. However, because of the high hardness of ceramic reinforcement, conventional turning and diamond machining, which represent the most widely used machining methods, currently provide the only options for machining metal matrix composites to the required accuracy. Even for the cutting wedges made of polycrystalline diamond and cubic boron nitride the tool wear is intense [3, 6]. Nevertheless, this technique involves low material removal rates and high tool wear, which result in high machining cost. Therefore, there is still a demand for hard to machine materials processing methods capable of enhancing high material removal rates, improving tool wear, and increasing the surface quality of the workpiece [7–9].

One way to improve the machinability of difficult to cut materials is to employ the thermal softening ability of a heat source to fluxing the material. The heat source is focused in front of a cutting tool to soften the workpiece material just ahead of the cutting tool thereby lowering the forces required to cut the material [1]. The technology is called a hybrid process of the machining, where the material is heated via laser irradiation and soft surface layer is machined by defined cutting edge of the wedges. This makes the material softer and easier to plastic deformation. In comparison with conventional machining this method significantly increases wedges durability [1, 7, 10–13]. LAM reduces the tool wear and cost of machining by reducing machine hours per part [14, 15]. In a study, Chang and Kuo [10] found a reduction of 20–22% in cutting force, which is an indicator of reducing the tool wear, and an increase in the profitability of using LAM during planning of alumina ceramics.

Research presented in this chapter builds on previous work by the author to further under‐ stand the process of laser-assisted turning. The present paper describes the experimental setup and characterization of a LAM process of a Si3N4 ceramics, cemented carbide layer, and metal matrix composites with a particular focus on the diamond tool wear and surface roughness.

#### **2. Experiments**

In the research, the standard metal matrix composite produced by the DURALCAN company was used as the workpiece material. The composite of a chemical composition similar to the AlSi9Mg alloy (9.2% Si, 0.6% Mg, 0.1% Fe) is reinforced with SiC particles with a particle size of between 8 to 15 μm and mass fraction about 20% (**Figure 1**). The sample has 50 mm in diameter and 10 mm length and before LAM it was painted by absorptive coat to improve absorbability of laser beam.

**Figure 1.** Microstructure of metal matrix composite used in research. The SiC particles are distributed through the ma‐ trix.

The machining tests were carried out on TUM 35D1 lathe with infinitely variable adjustment of rotational speed.

In order to determine the laser heating conditions for materials, there were materials that have been investigated, presented in **Table 1**.


**Table 1.** Characteristics of the samples.

Mechanical and physical properties like high thermal resistance, high hardness, good corrosive resistance and chemical stability have encouraged use of ceramics and SiC reinforced alumi‐ nium-based composites in several engineering applications. The brittle and hard nature of these ceramics makes them difficult to machine using conventional techniques and damage caused to the surface while machining affects efficiency of components. Laser-assisted machining (LAM) has recently emerged as a potential technique for attaining high material

MMCs and silicon nitride are increasingly attractive for engineering applications because of their superior combination of properties such as high strength, hardness, good corrosive resistance, and wear resistance. These materials are finding potential applications in the manufacture of automotive components such as some elements of cars, aerospace and medical [4, 5]. However, because of the high hardness of ceramic reinforcement, conventional turning and diamond machining, which represent the most widely used machining methods, currently provide the only options for machining metal matrix composites to the required accuracy. Even for the cutting wedges made of polycrystalline diamond and cubic boron nitride the tool wear is intense [3, 6]. Nevertheless, this technique involves low material removal rates and high tool wear, which result in high machining cost. Therefore, there is still a demand for hard to machine materials processing methods capable of enhancing high material removal rates,

One way to improve the machinability of difficult to cut materials is to employ the thermal softening ability of a heat source to fluxing the material. The heat source is focused in front of a cutting tool to soften the workpiece material just ahead of the cutting tool thereby lowering the forces required to cut the material [1]. The technology is called a hybrid process of the machining, where the material is heated via laser irradiation and soft surface layer is machined by defined cutting edge of the wedges. This makes the material softer and easier to plastic deformation. In comparison with conventional machining this method significantly increases wedges durability [1, 7, 10–13]. LAM reduces the tool wear and cost of machining by reducing machine hours per part [14, 15]. In a study, Chang and Kuo [10] found a reduction of 20–22% in cutting force, which is an indicator of reducing the tool wear, and an increase in the

Research presented in this chapter builds on previous work by the author to further under‐ stand the process of laser-assisted turning. The present paper describes the experimental setup and characterization of a LAM process of a Si3N4 ceramics, cemented carbide layer, and metal matrix composites with a particular focus on the diamond tool wear and surface roughness.

In the research, the standard metal matrix composite produced by the DURALCAN company was used as the workpiece material. The composite of a chemical composition similar to the AlSi9Mg alloy (9.2% Si, 0.6% Mg, 0.1% Fe) is reinforced with SiC particles with a particle size of between 8 to 15 μm and mass fraction about 20% (**Figure 1**). The sample has 50 mm in

improving tool wear, and increasing the surface quality of the workpiece [7–9].

profitability of using LAM during planning of alumina ceramics.

removal rates.

90 Diamond and Carbon Composites and Nanocomposites

**2. Experiments**

The example of laser-assisted turning of ceramics Si3N4 is illustrated in **Figure 2**.

The experimental setup applied for turning of Si3N4 ceramics is presented in **Figure 3**, while the setup for turning of MMC in LAM conditions is presented in **Figure 4**.

**Figure 2.** View of ceramics Si3N4 (Ø14×4.5×104mm) heated by laser beam.

**Figure 3.** The scheme of experimental set-up. Ceramics turning with LAM, designations: *A*—heating area by laser beam, *B*—zone of machining, *d*—workpiece diameter [1].

The 30 degree angle between heated area and zone of machining was selected **Figure 4**. The tools wear were measured for each trial of experiment. The surface roughness of the sample was measured by surface analyzer (Hommel T500). The machined surface finish was charac‐ terized by the average arithmetic roughness (R*a* parameter). Tool wear was measured on the primary flank face by the optical microscope.

The example of laser-assisted turning of ceramics Si3N4 is illustrated in **Figure 2**.

the setup for turning of MMC in LAM conditions is presented in **Figure 4**.

92 Diamond and Carbon Composites and Nanocomposites

**Figure 2.** View of ceramics Si3N4 (Ø14×4.5×104mm) heated by laser beam.

beam, *B*—zone of machining, *d*—workpiece diameter [1].

The experimental setup applied for turning of Si3N4 ceramics is presented in **Figure 3**, while

**Figure 3.** The scheme of experimental set-up. Ceramics turning with LAM, designations: *A*—heating area by laser

**Figure 4.** The scheme of experimental setup. (a) MMC laser-assisted turning, (b) Ceramics turning with LAM, designa‐ tions: *A*—heating area by laser beam, *B*—zone of machining, *d*—workpiece diameter, *dl* —laser beam diameter [1].

LAM was carried out with a 2.6 kW CO2 laser (2.6kW Trumpf, type TLF2600t).

The workstation for Si3N4 ceramics was equipped with remotely controlled temperature measurement system, which works on the base of infrared detection of irradiation. The measurement data logging gives an opportunity of on-line tracking temperatures in: *θ*0 machined workpiece temperature in laser beam power spot size on the heated surface; *θ*2 cutting tool point on the machined surface temperature **Figure 3**). The laser-assisted turning process was done in two steps. The ceramic bush with *n* rotational speed was heated in *to* period time up to reach the target of suitable workpiece temperature. The next step was the initiation of feed motion, and then the laser power spot size within wedge tool was straight moved. The constant temperature during laser-assisted turning process was ensured by dedicated computer program. The obtained *θ*<sup>2</sup> temperature data logging was compared with manual typing process temperature. The on-line program has changed laser beam power in case of appeared differences between those temperatures. The system power control was worked on the base of negative feedback.

The research was carried out for these cutting parameters:

Si3N4: cutting speed *vc* = 10 m/min, feed rate *f* = 0.04mm/rev, depth of cut *ap* = 0.05mm;

MMC:cutting speed *vc* = 10–100m/min, feed rate *f* = 0.04–0.1 mm/rev, depth of cut *ap* = 0.1mm;

Cemented carbide: cutting speed *vc* = 20 m/min, feed rate *f* = 0.04 mm/rev, laser power *P* = 600– 2600W.

The characteristics of applied edges are shown in **Table 2**.


**Table 2.** Characteristics of applied edges.

#### **3. Results**

#### **3.1. Laser surface heating**

In **Figure 5**, an example of two point measurements (*θ*0 and *θ*2) for Si3N4 ceramic is presented. The slow heating process of rotating Si3N4 ceramics up to required and accepted *θ*<sup>2</sup> temperature into control system was done in *t* = 0‐ to time interval only. The system control unit was increasing laser power irradiation after switching on the *vf* feed motion. The laser power irradiation and cutting tool were displaced; therefore, the laser power increase was necessary to get *θ*<sup>2</sup> constant temperature of the cutting tool point on the machined surface. As a result, the *θ*0 temperature increase was noticed in the first phase. The explanation of this phenomenon

**Figure 5.** Temperature distributions of the *θ* 0 and *θ* 2 during laser-assisted turning within temperature control.

is low thermal conductivity of heated ceramics; therefore, the first phase heating time during feed on required higher laser power irradiation level.

**Type of edge material Symbol of**

94 Diamond and Carbon Composites and Nanocomposites

3 Oxide ceramics (Al2O3 + ZrO2)

**3. Results**

**Table 2.** Characteristics of applied edges.

**3.1. Laser surface heating**

2 Sintered carbide H10S Cemented

4 Mixed ceramics (Al2O3 + TiCN) MC2 Cemented

5 Polycrystalline diamond PCD KD100 Cemented

increasing laser power irradiation after switching on the *vf*

**insert material**

1 Sintered carbide inserts with fine-grain substrate KC5510 (PVD) TiAlN SNMG120408

In **Figure 5**, an example of two point measurements (*θ*0 and *θ*2) for Si3N4 ceramic is presented. The slow heating process of rotating Si3N4 ceramics up to required and accepted *θ*<sup>2</sup> temperature into control system was done in *t* = 0‐ to time interval only. The system control unit was

irradiation and cutting tool were displaced; therefore, the laser power increase was necessary to get *θ*<sup>2</sup> constant temperature of the cutting tool point on the machined surface. As a result, the *θ*0 temperature increase was noticed in the first phase. The explanation of this phenomenon

**Figure 5.** Temperature distributions of the *θ* 0 and *θ* 2 during laser-assisted turning within temperature control.

**Coating Insert**

SNMG120408

SNGN 120708T 02020

SNGN 120708T 02020

TPGN110304F

feed motion. The laser power

Uncoated

Uncoated

Uncoated

Uncoated

AC5 Cemented

In **Figures 6** and **7** the courses of MMC surface temperature during heating with turning kinematics are shown. Temperature in heated area by laser beam succumb to stabilization (**Figure 7**), varying with the increase of temperature in time on *B* area (**Figures 6** and **7**). This is due to the heat accumulation in the examined sample. Range of temperature *Θ1* equals about 30°C (**Figure 6**), varying based on the different thickness of absorption layer (gouache) on sample surface as well as differences of surface texture for example surface roughness. There is noticeably varied increase of temperature *Θ1*from 223 to 510°C (**Figure 7**) in plan machining area (*B*).

**Figure 6.** Courses of temperature *Θ*1 (*B* zone) during heated MMC by laser beam with maximum and minimum values.

**Figure 7.** Exemplary courses of temperature *Θ1* in *A* (*Θ0*) and *B* (*Θ1*) areas during laser heating metal matrix composite.

The **Figure 8** shows the temperature course in time for cemented carbide sample, close to the place of the incidence of the laser beam for different laser powers with heating speed of 20 m/min. The courses of the temperature measurement are stable and do not increase over time.

**Figure 8.** The temperature profile over time for cemented carbide with different laser power in place of the heating area *vl* = 20 m/min.

#### **3.2. Tool wear**

In **Figure 9**, the investigation results are shown. The laser heated machined ceramic surface temperature increase has decreasing effect on the **VBc** wear insert tool indicator and the lowest values of it for *θ*2 = 1400–1500°C temperatures were noticed. The value of wear inserts tool in those temperatures is more than 5 times lower than wear without using laser heating of machined layer. The inserts tool abrasive and adhesive wear symptoms were observed in the range of investigated temperatures. However, thermal wear symptoms were not apparent. The *θ*<sup>2</sup> = 1400°C as an optimal temperature was selected for further investigation of machined surface laser heating, because the selection of *θ*0 > 2000°C results in an evaporation risk on machined material.

**Figure 9.** Polycrystalline diamond (PCD) inserts wear in function of the machined surface temperature after 1 tool pass.

In **Figure 10**, wear distributions of investigated tool wedges during laser-assisted turning of ceramic process for selected temperature are presented. The value of the tool insert wear indicator in the range of 1300 and 1400°C temperatures is diversified for the first few passes. However after the next few passes, the difference becomes smaller and after 10 minutes time of machining is insignificant. It could be expected that in real-time analysis the presented wear plots below will look a little bit different, because high temperature exposure of the tool wedges will be much longer than investigated time interval.

**Figure 10.** Polycrystalline diamond (PCD) wedges wear distributions during laser-assisted turning process for differ‐ ent Si3N4 machined surface temperature. The applied parameters: *vc* = 10 m/min, *ap* = 0.05 mm, *f* = 0.04 mm/rev.

The results of the influence of the heating temperature (*θ*2=1400°C) cut layer of ceramic Si3N4 on tool wear for different wedges are shown in **Figure 11**. It can be observed that the lowest value of tool wear was for PCD and the largest for ceramics wedges.

**Figure 11.** Tool wear for different cutting wedges.

20 m/min. The courses of the temperature measurement are stable and do not increase over

**Figure 8.** The temperature profile over time for cemented carbide with different laser power in place of the heating

In **Figure 9**, the investigation results are shown. The laser heated machined ceramic surface temperature increase has decreasing effect on the **VBc** wear insert tool indicator and the lowest values of it for *θ*2 = 1400–1500°C temperatures were noticed. The value of wear inserts tool in those temperatures is more than 5 times lower than wear without using laser heating of machined layer. The inserts tool abrasive and adhesive wear symptoms were observed in the range of investigated temperatures. However, thermal wear symptoms were not apparent. The *θ*<sup>2</sup> = 1400°C as an optimal temperature was selected for further investigation of machined surface laser heating, because the selection of *θ*0 > 2000°C results in an evaporation risk on

**Figure 9.** Polycrystalline diamond (PCD) inserts wear in function of the machined surface temperature after 1 tool

In **Figure 10**, wear distributions of investigated tool wedges during laser-assisted turning of ceramic process for selected temperature are presented. The value of the tool insert wear

time.

96 Diamond and Carbon Composites and Nanocomposites

area *vl*

= 20 m/min.

machined material.

pass.

**3.2. Tool wear**

Wedge wear of the AC5 in temperature 1300°C was minimum and at three times lower than at room temperature (**Figure 12**). Tool wear of ceramics as compared to the cemented carbide was very large and the wedges are classified as unable to withstand such difficult conditions.

**Figure 12.** Ceramics AC5 wedge wear in function of machining of the machined surface temperature Θ2.

**Figure 13** shows the influence of LAM on the flank wear during turning of Al-SiC metal matrix composites. These results indicated that as a consequence of heating the wedge wear was decreased significantly for tested tools. In **Figure 13**, it can be observed that for LAM, wedge wear is about 31% lower in comparison to conventional turning with cemented carbide and 35% for polycrystalline diamond wedge. However, it is equally important to note that when comparing all results, a laser power of 1000W is the most likely reason giving the optimum tool wear when machining Al/SiC metal matrix composite. The research confirmed the thesis that the softening of the matrix material composite allows pushing in or sliding hard reinforc‐ ing particles in the surface layer thus reducing the wear of cutting wedges.

**Figure 13.** The evolution of the tool wear VBc indicator during turning the metal matrix composites with different feed rate values, the diamond and cemented carbide wedges.

The main wear pattern was observed as regular flank wear. Flank wear is due to the abrasive action of the reinforced particles presence in the MMC. The hard particles SiC of hardness about 2600HV grind the flank face of the cutting tools in similar way as a grinding wheel during machining of materials.

#### **3.3. Machined surface roughness**

Wedge wear of the AC5 in temperature 1300°C was minimum and at three times lower than at room temperature (**Figure 12**). Tool wear of ceramics as compared to the cemented carbide was very large and the wedges are classified as unable to withstand such difficult conditions.

98 Diamond and Carbon Composites and Nanocomposites

**Figure 12.** Ceramics AC5 wedge wear in function of machining of the machined surface temperature Θ2.

ing particles in the surface layer thus reducing the wear of cutting wedges.

**Figure 13** shows the influence of LAM on the flank wear during turning of Al-SiC metal matrix composites. These results indicated that as a consequence of heating the wedge wear was decreased significantly for tested tools. In **Figure 13**, it can be observed that for LAM, wedge wear is about 31% lower in comparison to conventional turning with cemented carbide and 35% for polycrystalline diamond wedge. However, it is equally important to note that when comparing all results, a laser power of 1000W is the most likely reason giving the optimum tool wear when machining Al/SiC metal matrix composite. The research confirmed the thesis that the softening of the matrix material composite allows pushing in or sliding hard reinforc‐

**Figure 13.** The evolution of the tool wear VBc indicator during turning the metal matrix composites with different feed

rate values, the diamond and cemented carbide wedges.

One of the important indexes of finished product is surface quality. For these experiments, average machined surface roughness values of LAM were found to be lower than those of the conventional turning for all tested tools.

Machined surface roughness of Si3N4 ceramics after turning with polycrystalline diamond wedges (PCD) indicates some regularities (**Figure 14**). The increase of temperature in cutting layer (in majority of reported cases) increases the height of roughness profile. The increase of tool wear at the different cutting temperatures does not change this regularity. Therefore a thesis can be stated, that microroughness surface height is not related to the cutter's microge‐ ometry, but temperature of cutting layer in laser-assisted process.

Similar to courses described earlier were those obtained for the turning with cemented carbide wedges KC5510 (**Figure 16**).

**Figure 14.** Machined surface roughness in function of machined Si3N4 surface temperature, after turning with PCD pol‐ ycrystalline diamond wedges.

(for *ts*≈ 2 min) *Si3N4*ceramic machined surface, it can be noticed that in the analyzed temperature range of cutting layer, surface roughness after turning with carbide KC5510 cutter is higher than that generated for the PCD wedge (**Figure 15**).

**Figure 15.** The initial roughness (ts ≈ 2 min) of machined surface for difference temperatures of machining layer after PCD and KC5510 wedges turning.

The increase of the machined surface roughness of Si3N4 ceramics, together with the temper‐ ature growth, can be explained by changes in the decohesion mechanism of the workpiece. Lower temperatures of the process are dominated by a brittle cracking, while the higher ones are dominated by a plastic flow of work material. Major role in the creation formation of machined surface during turning with a laser heating of a cutting layer have microchips with tendency to adhere to workpiece, and thus to increase in the height of microroughness.

**Figure 16.** Machined surface roughness distributions during laser-assisted turning process for different machined sur‐ face Si3N4 temperature, after turning with KC5510 carbide wedges.

The higher temperature of ceramic materials in machining layer during the LAM process improves the cutting ability evaluated by a tool wear [11] but has an adverse effect on quality of machined surface.

Average width of the profile roughness grooves (*RSm*) are, in the investigated range, higher than the feed rate values (**Figure 17**). This indicates difficulties in formation mechanism of the machined surface, inducing serious disturbances in kinematic-geometric projection of the cutter into the workpiece.

**Figure 15.** The initial roughness (ts ≈ 2 min) of machined surface for difference temperatures of machining layer after

The increase of the machined surface roughness of Si3N4 ceramics, together with the temper‐ ature growth, can be explained by changes in the decohesion mechanism of the workpiece. Lower temperatures of the process are dominated by a brittle cracking, while the higher ones are dominated by a plastic flow of work material. Major role in the creation formation of machined surface during turning with a laser heating of a cutting layer have microchips with tendency to adhere to workpiece, and thus to increase in the height of microroughness.

**Figure 16.** Machined surface roughness distributions during laser-assisted turning process for different machined sur‐

face Si3N4 temperature, after turning with KC5510 carbide wedges.

PCD and KC5510 wedges turning.

100 Diamond and Carbon Composites and Nanocomposites

**Figure 17.** Machined surface roughness *RSm* parameter distributions for the different machined Si3N4 surface tempera‐ ture, after turning with PCD wedges.

**Figure 18** shows that surface texture has a random character without any wedge mapping in the investigated feed range (f = 0.04 mm/rev). However, the graph of 3D profile (**Figure 18a**) and power spectral density (PSD) graph reveal that dominant peak is different from the applied feed rate (**Figure 18b**).

**Figure 18.** View of laser-assisted turning Si3N4 ceramics by polycrystalline diamond (PCD) wedges. (a) surface texture and (b) power spectrum density of the roughness profile. The applied parameters: *θ*<sup>2</sup> = 1400°C, *vc* = 10 m/min, *ap* = 0.05 mm, *f* = 0.04 mm/rev.

Nevertheless, in case of turning with KC5510, clear microstructure's orientation can be seen (**Figure 19**), which is attributed to the kinematic-geometric projection of the cutter into the workpiece. This observation is confirmed by a PSD diagram (**Figure 19b**), where the modal value of profile component has a frequency close to the feed rate.

**Figure 19.** View of Si3N4 ceramics surface texture (a) and power spectrum density of the roughness profile (b) after laser-assisted turning within carbide (KC5510) wedges. Parameters: *θ*2 = 1400°C, *vc* = 10m/min, *ap* = 0.05mm, *f* = 0.04mm/rev.

The machined surface roughness is one of the most important quality indicators of the MMC materials. **Figure 20** shows the influence of LAM on the machined surface roughness *Ra* parameter, during turning of Al/SiC metal matrix composite with different cutting wedges. Significant differences in the roughness of machined surface were observed for the laserassisted and conventional turning (*P*= 0 W) with the carbide wedges. However, for the conventional turning, it is seen that average machined surface roughness value is higher than that obtained for the LAM. It could be explained by the fact that, in conventional turning, SiC particles were pushed through the soft matrix or pushed out of matrix during machining. Consequently, cracks and pits were formed on the machined surface (**Figure 10**), which are inducing poor surface finish. During hybrid machining, laser heated layer is being cut in the liquid or semi-liquid state, so that fills the grooves in the machined surface generated after plowing or crumbling out the reinforcement particles. This mechanism reduces the average height of machined surface roughness.

**Figure 20.** The influence of a laser's beam power on machined surface roughness *Ra* after LAM and conventional turn‐ ing of aluminum-ceramic via different cutting wedges material.

The analysis of power spectrum density (PSD) of the roughness profile revealed that machined surface texture has random character with the significant disturbances of the cutting edge kinematic-geometric projection. The PSD chart has dominant component different from the applied feed rate. The analysis of power spectrum density for different values of feeds (*f1* = 0.07 mm/rev, *f2* = 0.1 mm/rev) reveals significant disturbances of the cutting edge kinematicgeometric projection (**Figure 21**).

Nevertheless, in case of turning with KC5510, clear microstructure's orientation can be seen (**Figure 19**), which is attributed to the kinematic-geometric projection of the cutter into the workpiece. This observation is confirmed by a PSD diagram (**Figure 19b**), where the modal

**Figure 19.** View of Si3N4 ceramics surface texture (a) and power spectrum density of the roughness profile (b) after laser-assisted turning within carbide (KC5510) wedges. Parameters: *θ*2 = 1400°C, *vc* = 10m/min, *ap* = 0.05mm, *f* =

The machined surface roughness is one of the most important quality indicators of the MMC materials. **Figure 20** shows the influence of LAM on the machined surface roughness *Ra* parameter, during turning of Al/SiC metal matrix composite with different cutting wedges. Significant differences in the roughness of machined surface were observed for the laserassisted and conventional turning (*P*= 0 W) with the carbide wedges. However, for the conventional turning, it is seen that average machined surface roughness value is higher than that obtained for the LAM. It could be explained by the fact that, in conventional turning, SiC particles were pushed through the soft matrix or pushed out of matrix during machining. Consequently, cracks and pits were formed on the machined surface (**Figure 10**), which are inducing poor surface finish. During hybrid machining, laser heated layer is being cut in the liquid or semi-liquid state, so that fills the grooves in the machined surface generated after plowing or crumbling out the reinforcement particles. This mechanism reduces the average

**Figure 20.** The influence of a laser's beam power on machined surface roughness *Ra* after LAM and conventional turn‐

value of profile component has a frequency close to the feed rate.

102 Diamond and Carbon Composites and Nanocomposites

0.04mm/rev.

height of machined surface roughness.

ing of aluminum-ceramic via different cutting wedges material.

In **Figure 21b**, the orientation of surface texture was observed as a result of turning kinematics; however, a power spectrum density diagram did not reveal a component related to the employed feed rate. **Figure 21a** and **21b** shows a cavity created after crumbling out the SiC particles from metal matrix.

**Figure 21.** 3D topographic images (machined surface of metal matrix composite and power spectrum density) of sur‐ face roughness profiles after conventional turning of non-coated carbide H10S wedge for feed rate: (a) *f* = 0.07 mm/rev, (b) *f* = 0.1 mm/rev.

However, in the case of conventional turning, there was an increase in average surface roughness for analyzed sample in comparison with LAM. The formation of grooves on the sample as a result of moving of SiC particles by cutting tool is a major cause of the deterioration of surface roughness with traditional turning. With a LAM the temperature in cutting zone increased and leads to weakening the connections between the SiC particles and aluminium matrix. In most cases, it causes pushing the particle outside or in to the aluminium alloy. The same observation had Altinkok [16] in research.

**Figure 22** shows a comparison of the surface profiles of the machined surface by conventional cutting and LAM. The profiles show the difference of surface finish achieved.

**Figure 22.** Comparison of machined surface roughness profiles for the same cutting conditions. Laser Assisted Machin‐ ing (a) and conventional turning (b)

#### **4. Conclusions**

Laser enhanced conventional turning to heat the cutting materials, as well as to change the microstructure or locally harden the material in front of the cutting tool. This process is carried out in order to facilitate the machining due to its softening and change of the workpiece's deformation behavior. The local temperature of the material in the shear deformation zone plays an important role in the thermally enhanced machining process.

The investigations on laser-assisted turning of difficult to cut materials confirmed that laser heating of the machined surface causes significant improvement in turning process, which is estimated by tool wear and surface roughness indicators.

The low tool wear intensity in a suitable temperature is crucial for increasing the tool life. This opportunity enables the significant reduction of cost tool production.

The flank face wear was observed during laser-assisted and conventional turning. The abrasive tool wear mechanism is dominant in turning of these materials. Turning with laser heating improves machined surface roughness in comparison with conventional turning for all examined wedges.

#### **Acknowledgements**

However, in the case of conventional turning, there was an increase in average surface roughness for analyzed sample in comparison with LAM. The formation of grooves on the sample as a result of moving of SiC particles by cutting tool is a major cause of the deterioration of surface roughness with traditional turning. With a LAM the temperature in cutting zone increased and leads to weakening the connections between the SiC particles and aluminium matrix. In most cases, it causes pushing the particle outside or in to the aluminium alloy. The

**Figure 22** shows a comparison of the surface profiles of the machined surface by conventional

**Figure 22.** Comparison of machined surface roughness profiles for the same cutting conditions. Laser Assisted Machin‐

Laser enhanced conventional turning to heat the cutting materials, as well as to change the microstructure or locally harden the material in front of the cutting tool. This process is carried out in order to facilitate the machining due to its softening and change of the workpiece's deformation behavior. The local temperature of the material in the shear deformation zone

The investigations on laser-assisted turning of difficult to cut materials confirmed that laser heating of the machined surface causes significant improvement in turning process, which is

plays an important role in the thermally enhanced machining process.

estimated by tool wear and surface roughness indicators.

cutting and LAM. The profiles show the difference of surface finish achieved.

same observation had Altinkok [16] in research.

104 Diamond and Carbon Composites and Nanocomposites

ing (a) and conventional turning (b)

**4. Conclusions**

The author wishes to gratefully acknowledge Mr. Marian Jakowiak for valuable advice, support, time, and available materials that contributed in creating the work. The part of presented research results, executed under the domestic project LIDER of No 141/L-5/2013, was funded with grants for education allocated by the National Centre for Research and Development

#### **Author details**

Damian Przestacki

Address all correspondence to: damian.przestacki@put.poznan.pl

Poznan University of Technology, Poznan, Poland

#### **References**

