REVIEW
Biomedical applications of nanotechnology
Ana P. Ramos
1
& Marcos A. E. Cruz
1
& Camila B. Tovani
1
& Pietro Ciancaglini
1
Received: 28 November 2016 /Accepted: 26 December 2016 /Published online: 13 January 2017
#
International Union for Pure and Applied Biophysics (IUPAB) and Springer-Verlag Berlin Heidelberg 2017
Abstract The ability to investigate substances at the molecu-
lar level has boosted the search for materials with outstanding
properties for use in medicine. The application of these novel
materials has generated the new research field of
nanobiotechnology, which plays a central role in disease di-
agnosis, drug design and delivery, and implants. In this re-
view, we provide an overview of the use of metallic and metal
oxide nanoparticles, carbon-nanotubes, liposomes, and
nanopatterned flat surfaces for specific biomedical applica-
tions. The chemical and physical properties of the surface of
these materials allow their use in diagnosis, biosensing and
bioimaging devices, drug delivery systems, and bone substi-
tute implants. The toxicology of these particles is also
discussed in the light of a new field referred to as
nanotoxicology that studies the surface effects emerging from
nanostructured materials.
Keywords Nanobiotechnology
.
Inorganic particles
.
Liposomes
.
Nanopatterned surfaces
Introduction
Nanotechnology has many definitions and applications.
However, all definitions highlight the design and development
of highly ordered bottom–up nanostructured materials that
offer specific responses when exposed to certain stimuli
(Saji et al. 2010). Surface chemistry and physics Btune^ the
applications of nanosized materials. The concentration of
atoms on the surface of these systems represents up to 90%
of their total mass and results in enhanced reactivity. In this
sense, modifying the surface of a nanomaterial in different
ways can produce materials with distinct biological properties
and functionalities for a specific end applica tion and with
improved solubility under physiological conditions (Gupta
et al. 2007;Kasemo2002).
Nanotechnology products have become increasingly useful
in biomedicine and have led to the advent of a hybrid science
named nanobiotechnology (Saji et al. 2010). Nanomaterials
have noteworthy applications in nanobiotechnology, particu-
larly in diagnosis, drug delivery systems (Faraji and Wipf
2009), prostheses, and implants. Nanoscale materials integrate
well into biomedical devices because most biological systems
are also nanosized. The materials commonly used to develop
these nanotechnology products are inorganic and metal nano-
particles, carbon nanotubes, liposomes, and metallic surfaces
(Liu et al. 2016a). By using chemical or physical methods and
taking advantage of specific biological reactions, such as the
antibody–antigen interaction, receptor–ligand interaction, and
DNA–DNA hybridization, it is possible to conjugate
biospecific molecules with nanoparticles. Surface chemistry
(composition) (Castner and Ratner 2002; Moyano and
Rotello 2011), surface physics (topography and roughness)
(McNamara et al. 2010; Yim et al. 2010), surface thermody-
namics (wettability and free energy) (Menzies and Jones
2010), and their toxicological effects determine the specific
application of nanomaterials.
In this review we discuss the biomedical applications of
nanoparticles and nanopatterned surfaces (Fig. 1), including
surface features and modifications which are responsible for
tuning their response when in contact with biological environ-
ments. The design of nanostructures by controlling their
* Ana P. Ramos
anapr@ffclrp.usp.br
1
Departamento de Química, Faculdade de Filosofia, Ciências e Letras
de Ribeirão Preto (FFCLRP), Universidade de São Paulo (USP),
14040-901 Ribeirão Preto, SP, Brazil
Biophys Rev (2017) 9:79–89
DOI 10.1007/s12551-016-0246-2
surface properties is presented as a strategy to achieve im-
proved responses aimed at a specific application. In this con-
text we focus on the use of inorganic (metallic and metal
oxide) and organic (carbon-nanotubes and liposomes) nano-
particles and nanopatterned flat surfaces in diagnosis, biosens-
ing and bioimaging devices, drug delivery systems, and bone-
substituting implants. The toxicology of these particles is also
discussed in the light of a new field referred to as
nanotoxicology that studies the surface effects emerging from
nanostructured materials.
Biomedical applications of metal oxide nanoparticles
Metal oxide nanoparticles have been employed to construct
several medical devices. The magnetic properties of iron oxide
have been used for therapeutic and diagnostic purposes, such as
contrast agents for magnetic resonance imaging, magnetic par-
ticle imaging, and ultrasonic techniques (e.g. magneto-motive
ultrasound (Oh et al. 2006), photoaco ustic imaging, and mag-
netic particle hyperthermia (Gupta and Gupta 2005;Liuetal.
2016b). The electronic structure of zinc oxide (ZnO) is useful
for biomedical applications; for example, the intrinsic fluores-
cence of ZnO nanowires has been employed to image cancer
cells (Hong et al. 2011). To this end, functionalization of the
surface of ZnO nanowires increases their solubility in water and
their biocompatibility and reduces their c ellular toxicity.
Functionalization of the ZnO surface with specific biomole-
cules creates photosensitive biosensors (Liu et al. 2006).
The high surface area of nanoparticles favors the prompt
adsorption of plasma proteins (Deng et al. 2009). Hence, the
chemical composition and physical topography of the surface
of nanoparticles and the combination of these properties (wet-
tability, surface-free energy) tailor the interaction of the parti-
cles with different compounds and dictate their end applica-
tion (Fig. 2).
T itanium oxid e (T iO
2
) has a wide range of biomedical appli-
cations (Fei Y in et al. 2013). For instance, in bone-substituting
materials, the biofluid first makes contact with a thin TiO
2
layer
that spontaneously emerges on the top surface of metallic titani-
um (Hanawa 201 1 ;Fengetal.2003). This has motivated the use
of T iO
2
nanoparticles for bone regeneration (He et al. 2008;
Kubota et al. 2004; Brammer et al. 2012; Tan et al. 2012).
Zirconium oxide has recently been used for dental implants
because, like titanium, it is compatible with the same type of
hard tissues (Koch et al. 2010; Özkurt and Kazazoğlu 201 1).
Metal nanoparticles
The strong optical absorption related to the surface plasmon
resonance of noble metals makes them suitable for construct-
ing molecular contrast devices (Liao et al. 2006; Bhattacharya
and Mukherjee 2008). Absorption and scattering in the visible
and near-infrared regions have stimulated the application of
materials containing metal nanoparticles in the fields of sens-
ing and diagnosis. Gold nanoparticles can be deposited on
appropriate substrates or added to substrate formulation to
enhance luminescence (Bhattacharya and Mukherjee 2008).
The application of this technology depends on the size and
geometry of the particles because these determine their
absorption/scattering properties. Gold nanorods absorb in the
near-infrared and have been used to monitor the blood flow
in vivo using photoacoustic imaging (Wang et al. 2005). The
literature contains examples of applications that used gold
nanocages, nanoshells, and nanospheres (Liao et al. 2006). It
is possible to modify the surface of gold nanoparticles with
sulfur-containing compounds because gold and sulfur have a
high chemical affinity (Schmidt and Healy 2009;Moyanoand
Rotello 2011). Modification of gold nanoparticles with
biospecific compounds enhances binding to specific tissues
(Faraji and Wipf 2009). For example, surface-labeled gold
Fig. 1 Nanobiotechnology and its main tools
Fig. 2 Correlation between the main chemical and physical properties of the surface of nanoparticles and the end nanobiotechnological application of
nanomaterials
80 Biophys Rev (2017) 9:79–89
nanoshells hav e been used to target can cer cells in vitro
(Bhattacharya and Mukherjee 2008) and the results confirmed
by optical microscopy.
Given all the advantages of using noble metal nanoparticles
for biomedical applications, silver nanoparticles (AgNPs)
have also attracted interest. Some biosensing applications are
based on spectral modifications due to aggregated particles
(Liao et al. 2006;MoyanoandRotello2011). These particles
display well-known antibacterial activity as well as an anti-
inflammatory action (Chaloupka et al. 2010). The preparation
of AgNPs is straightforward (Guidelli et al. 2011, 2012)as
their size is easy to control, and they can be incorporated into
different materials without difficulty (Guidelli et al. 2013;
Guidelli et al. 2016). Published reports describe the use of
silver as a coating in materials for cardiovascular implants
and central venous and neurosurgical catheters (Chaloupka
et al. 2010; Chen and Schluesener 2008).
Polymers used as bone cement (Alt et al. 2004) and wound
dressing (Tian et al. 2007) have also been loaded with AgNPs
and their antimicrobial responses compared to standard anti-
biotics and silver salts compared. Only AgNPs-bone cement
combines high antibacterial activity with low cytotoxicity as
compared to silver salt and gentamicin-bone cement (Alt et al.
2004). Controlled delivery coupled to small particle size is one
strategy to reduce toxicity. AgNPs have been added to latex
membranes used for skin regeneration (Guidelli et al. 2013),
with the membranes acting as a biomaterial and controlling
the nanoparticle delivery rate (Abukabda et al. 2016).
Carbon nanotubes
The physical and chemical properties of carbon nanotubes
(CNTs) have motivated their application in several areas of
science. Modification of the surface of these particles and their
functionalization with biological molecules at the molecular
level has increased their use in nanobiotechnology (Yang et al.
2007; Prato et al. 2008; Sharma et al. 2016). These modified
particles provide well-dispersed samples that are compatible
with physiological conditions (Williams et al. 2002). In this
context, nanotubes might be useful drug delivery vehicl es
because their nanometer size enables them to move easily
inside the body (Pastorin et al. 2006; Faraji and Wipf 2009).
The bioavailability of methotrexate, a drug used in cancer
therapy, increases when it is administered after being
immobilized on a double-functionalized carbon nanotube sur-
face (Pastorin et al. 2006). The active compound can be
inserted inside the tube or it can bind to the surface of the
particle with the aim to target and alter cell behavior at the
subcellular or molecular level. Moreover, biofunctionalized
single- or multi-walled CNTs can be taken up by a wide range
of cells, traffic through different cellular barriers (Yang et al.
2007), and interact with DNA (Pantarotto et al. 2004a).
Cationic CNTs complexed with plasmid DNA. can also be
taken up by cells and interact with DNA (Pantarotto et al.
2004b). CNTs are suitable scaffolds for the proliferation of
osteoblasts (Zanello et al. 2006; Zancanela et al. 2016)and
for the regeneration of bones (Zhao et al. 2010;Yoonetal.
2014). We have demonstrated that the toxicity of unmodified
single- and multi-walled CNTs is concentration-dependent
and that at concentrations of up to 10
−2
mg/mL, they particles
can be safely used in osteoblast cultures (Zancanela et al.
2016). We recently proposed the use of collagen-modified
calcium carbonate nanotubes as a new generation of tubular
structures for bone regeneration (Tovani et al. 2016).
Liposomes and nanobiotechnology
Liposomes are small artificial lipid-bilayer spherical vesicles
that were first reported by Bangham and Horne (1964).
Liposomes with different properties can be achieved by tuning
their composition, surface charge, and size. The rigidity and
fluidity of the bilayer can also be tailored by choosing specific
lipids (Akbarzadeh et al. 2013). These artificial membrane
models can be classified on the basis of their diameter. Small
unilamellar vesicles range in size from 20 to 100 nm, whereas
large unilamellar vesicles (LUVs) range from 200 to 1000 nm.
The vesicles consist of a single lipid bilayer and an internal
aqueous cavity. Liposomes can also be classified according to
the number of lipid bilayers. Multilamellar vesicles are formed
by multiple, concentric phospholipid bilayers intercalated
with aqueous compartments, with diameters ranging from
400 to 3500 nm (Akbarzadeh et al. 2013; Bilia et al. 2014).
Over the last two decades, liposomes have been widely
employed as drug delivery systems for cancer and gene ther-
apy and vaccines, among other uses (Madni et al. 2014;Liu
and Chen 2015). These vesicles can deliver a range of bioac-
tive compounds, such as antioxidants, antimicrobials, an d
angenic proteins (Simão et al. 2015). The functional ity of
these molecules is preserved after encapsulation (Benech
et al. 2002; Shehata et al. 2008; Akbarzadeh et al. 2013;
Gao et al. 2014). Moreover, compounds with different solu-
bility can be encapsulated inside the aqueous cavity or at the
surface of the lipid bilayers (Ghalandarlaki et al. 2014 ;
Reimhult 2015). Because liposomes are potentially atoxic,
degradable under physiological conditions, and non-immuno-
genic, they can be expected to deliver drugs with a low deg-
radation rate, with diminished collateral effects (Ravi-Kumar
2000; Ghalandarlaki et al. 2014).
More than materials for drug delivery, liposomes can also
be used as biomimetic models to study how membranes inter-
act with hydrophobic drugs and proteins. Actually, protein-
associated liposomes (proteoliposomes) can be employed to
investigate the action of photosensitive dyes applied during
photodynamic therapy (Bolfarini et al. 2012; Longo et al.
Biophys Rev (2017) 9:79–89 81
2012; Faria et al. 2015) to prevent/treat several diseases
(Daghastanli et al. 2004; De Lima Santos et al. 2005;
Zucolotto et al. 2007; Ciancaglini et al. 2012; Simão et al.
2015). Proteoliposomes can be used to study how the pro-
tein–lipid interactions specifically affect the surface of the
membrane model. The scientific community has highlighted
this approach as an excellent tool to understand biochemical
and biophysical phenomena and to evaluate biotechnological
applications (Camolezi et al. 2002; Ierardi et al. 2002 ;
Daghastanli et al. 2004; De Lima Santos et al. 2005;Rigos
et al. 2008;Santosetal.2009; Bolean et al. 2010, 2011, 2015;
Simão et al. 2010a, 2010b; Barbosa et al. 2011; Yoneda et al.
2013, 2014; Simão et al. 2015;Dongetal.2016;Elkhodiry
et al. 2016).
Nanotechnology to engineer the surface of metallic
implants
Nanotechnology has also found applications in tissue and im-
plant engineering. The possibility to enhance the surface area
of the material and to tune the roughness of its surface at the
nanometric scale should yield better biological responses of
osteogenic cells and effective mechanical contact between tis-
sue and implant. Titanium and its alloys are considered to be
the most attractive materials for bone replacement applications
(Rack and Qazi 2006). The widespread use of this metal is due
to its improved mechanical properties, high resistance to cor-
rosion, low surface reactivity, and acceptable biocompatibility
in vivo and in vitro. The live tissue heals in close apposition to
the metal, although there may be a thin fibrous layer separat-
ing the metallic implant and the bone that represents a failure
in the osteointegration process. In this context, it is necessary
to modify the surface of the implant to create a stronger bone–
implant interface and to achieve successful osteointegration
(Puleo and Nanci 1999; Le Guéhennec et al. 2007). At first
glance, modification of the surface should only change the
topography (Le Guéhennec et al. 2007). However, the addi-
tion of bioactive compounds and the creation of roughness at
the nanometer level appear to be more promising strategies for
biomedical applications. In the following subsections, we re-
view which factors must be modified on the surface of im-
plants if better host tissue responses are to be achieved. We
also describe how nanotechnology is being used to engineer
the surface of implants at the nanometer scale.
Why modify an implant surface?
The aim of modifying metallic surfaces is to improve the
contact between the implant and the live tissue in bone-
substitution applications (Jalota et al. 2007). Successful ortho-
pedic implant osteointegration relies on the quick and efficient
formation of bone tissue at the surface of an implant
(Albrektsson et al. 1981). A cell never encounters a complete-
ly clean surface; instead, it comes into contact with a surface
conditioned by water molecules, ions, and adsorbed proteins.
Therefore, water interactions, protein adsorption, and cell at-
tachment are the first events taking place at the tissue–implant
interface after implantation (Puleo and Nanci 1999). This con-
ditioned surface dictates cell attachment and the resulting mor-
phology and behavior of the cell. All these early events at the
bone–implant interface are determined by properties of the
surface, such as topography, wettability, charge, and chemical
composition (Chen et al. 2014). Surface engineering ensures
that an implant with an optimized surface is achieved by ma-
nipulating these properties to maximize anchorage of the im-
plant to the tissue. The need to modify the surface is clear from
clinical observations indicating that bone growth rate is higher
moving away from the implant surface than toward the im-
plant surface (Puleo and Nanci 1999).
In the case of bone-substitution materials, roughness gov-
erns the amount of bone tissue that is in close contact with the
surface of the implant (Wennerberg et al. 1998). In this con-
text, topographic features, such as valleys, peaks, and grooves,
act as points for cell anchorage and protein adsor ption
(Lampin et al. 1997; Anselme et al. 2000; Webster et al.
2000; Jayaraman et al. 2004). For materials with the same
surface chemistry, cell growth will be driven by topographic
features (Rosales-Leal et al. 2010). Rougher surfaces promote
higher adhesion of osteoblasts (Martin et al. 1995; Webster
and Ejiofor 2004) and dictate the metabolism of these cells
regulating gene expression (Brett et al. 2004), the synthesis of
collagen (Boyan et al. 2001; Wennerberg and Albrektsson
2009), and the activation of integrins (Khang et al. 2012).
Manipulation of the topography of a surface at the nanoscale
level has been shown to positively affect cell behavior (Khang
et al. 2012). Osteogenesis starts faster on surfaces organized at
the nanometer level than on smooth surfaces (Riehle et al.
2003;
Webster and Ejiofor 2004;Satoetal.2005).
Therefore, manipulation of the nanotopography of a biomate-
rial can stimulate and control cellular behavior such as attach-
ment, migration, spreading, gene expression, proliferation,
differentiation, and secretion of matrix components (Klymov
et al. 2013). Creating materials organized at the nanometric
level can be an effective strategy to target the cell recognition
process (Brunetti et al. 2010) and may affect the interaction of
solvent molecules with the surface, thus impacting the inter-
facial energy of the material (Kuna et al. 2009).
Wettability determines how cells and fluids interact with
surfaces. Wettability refers to the ability of a fluid to spread
on a given surface. It is related to the equilibrium of forces
acting at the solid–liquid interface and is governed by the to-
pography of the surface (Quéré 2008). In biomaterials science,
wettability is assessed by measuring contact angles (θ) between
a liquid drop and the surface (Menzies and Jones 2010). When
this liquid i s water, surfaces where the water droplets
82 Biophys Rev (2017) 9:79–89
spontaneously spread over the surface (θ < 90°) are considered
to be hydrophilic; if θ is >90°, the surface is considered to be
hydrophobic. However , this classical limit between a hydro-
philic and a hydrophobic surface has been reviewed due to
the specific structure of water molecules at the interface (Berg
et al. 1994;Vogler1998). The value of θ is measured at the
solid–liquid–gas interface and is univocally fixed by the chem-
ical nature of the different phases and the equilibrium forces
acting among these phases. Contact angles are mathematically
correlated by the Young–Dupré equation (Kwok and Neumann
1999), namely, cos(θ)=(γ
SG
− γ
SL
)γ
LG
,whereγ is the interfa-
cial tension between the solid (S), liquid (L), and gas (G)
phases. In particular, γ
SG
is called the surface free energy
(SFE) of a solid. SFE is an important parameter because it
can be determined by using chemical models that depend not
only on γ
SL
and γ
LG
, but also on spec ific intermolecular forces.
Hence,thetotalSFEcanbeseentobetheresultofthecombi-
nation of dispersive forces (γ
SG
d
)andpolarforces(γ
SG
p
) (Kwok
and Neumann 1999). Thus, the SFE of a biomaterial selectively
determines how either the polar or the non-polar portion of
proteins and cell membranes interact with the surfaces.
To achieve a specific effect for biomedical applications, the
type of p rotein and the manner in which the protein is
adsorbed onto the implant are more important than the amount
of adsorbed protein. Replacement of the initially adsorbed
proteins with other proteins that d isplay a higher affinity
dictates the bioactivity of the surface of a biomaterial (Tirrell
et al. 2002). The protein layer absorbed onto the surface of the
implant will determine how cells attach and spread, thereby
influencing cell maturation. Arima and Iwata (2007)showed
that albumin can block cell–implant interaction by strongly
binding to hydrophobic surfaces, thus inhibiting replacement
by other extra-cellular matrix (ECM) proteins. On the other
hand, when adsorbed albumin binds to hydrophilic groups, it
can indeed be replaced with ECM proteins. Consequently, the
type of protein and the type of binding govern cellular
adhesion and migration processes (Arima and Iwata 2007).
Cells can non-specifically adhere to surfaces through ionic
and van der Waal’s forces, or they can specifically adhere to
the surface via adsorbed protein clusters (Anselme 2000). In
osteoblasts, integrin recep tors recognize motifs such as
Arg-Gly-Asp (RGD) on protei ns like fib ronectin and
vitronectin to form local focal adhesions. The formation of
these adhesions activates a cascade of intracellular signaling
pathways that affect cell behavior (Hendesi et al. 2015).
Recognition of this integrin-mediated cell attachment
mechanism has inspired scientists to modify metallic surfaces
by immobilizing sequences of RGD-peptides (Ferris et al.
1999; Rammelt et al. 2006). Thus, surface modification adds
another dimension to bioactive coating by altering not only
the chemistry but also the topography; in this context,
manipulating t he SFE of the surface might be the more
promising strategy to stimulate the cell adhesion process.
Kilpadi et al. (2001) observed that hydroxyapatite (HAp)
adsorbs more ECM proteins and binds more integrins and
osteoblast precursor cells than pure titanium or steel.
Therefore, the addition of bioactive compounds (e.g., bioactive
polar groups) to the metallic surface could be an alternative
approach to tailoring SFE and cell behavior. On hydrophobic
surfaces, human fetal osteoblasts express significantly lower
levels of the α5andβ3 integrin subunits than cells cultured
on hydrophilic surfaces (Lim et al. 2005). Moreover , surfaces
containing polar groups (COOH and NH
2
) display an enhanced
activity of integrins, which leads to a higher adhesion and
spreading of fibroblast cells (Faucheux et al. 2004). Using
macrocrophage cultures, Hotchkiss et al. (2016) observed that
materials with high surface wettability produced a n anti-
inflammatory microenvironment through activation of macro-
phages and the production of cytokines, indicating it is crucial
to control wettability when attempting to improve the healing
response to biomaterials.
Bio-inspired modifications of surfaces
The addition of bioactive minerals inspired by the bone struc-
ture has been one of the most commonly used strategies to
modify metallic surfaces of the implant. Biomimetics is a de-
sirable strategy because it predefines nanochemical and/or
nanophysical structures. The manufacture of coatings based
on calcium phosphates (CaP) is common in biomaterials sci-
ence, and the application of such coatings to the implant sig-
nificantly affects the bone regeneration process (Surmenev
et al. 2014). HAp is a calcium orthophosphate mineral that
resembles the biological apatite found in bone tissue, where
HAp crystals are hierarchically organized into the array of
collagen fibers (Olszta et al. 2007). At the commercial level,
plasma spray (Surmenev 2011) and sputtered coating (Yang
et al. 2005) techniques produce apatite coatings. However,
such methods are complex and require extremely high tem-
peratures as well as expensive equipment. Moreover, these
mechanically based methods pose some challenges: homoge-
neous thickness and crystallinity are difficult to achieve, coat-
ing adhesion is low, CaP phases change during the coating
process, and particles are released from the surface (Le
Guéhennec et al. 2007; Surmenev 2011).
Inspired by the CaP growth process in in vivo systems,
physiological solutions or media have also been used to pro-
duce CaP coatings. This methodology allows the formation of
continuous CaP coatings with controlled surface topography
(Costa et al. 2012), but it can require long exposure times (Tas
2014). Simulated body fluid (SBF) is one of the most frequent-
ly used physiological solution (Kokubo et al. 1990; Cüneyt Tas
2000). SBF consists of a supersaturated CaP solution that sim-
ulates the pH and ionic composition of human body fluid and
is a standard method employed to evaluate the bioactivity of
materials (Kokubo and Takadama 2006). The use of
Biophys Rev (2017) 9:79–89 83