Construction and Biological Testing of Artificial Implants for ...

Construction and Biological Testing of Artificial

Implants for Peripheral Nerve Repair

Von der Fakultät für Mathematik, Informatik und Naturwissenschaften der

RWTH Aachen University zur Erlangung des akademischen Grades eines

Doktors der Naturwissenschaften genehmigte Dissertation

vorgelegt von

Diplom-Biologe

Andreas Kriebel

aus Leverkusen

Berichter: Prof. Dr. Jörg Mey

Prof. Dr. Hermann Wagner

Tag der mündlichen Prüfung: 26.11.2015

Diese Dissertation ist auf den Internetseiten der Universitätsbibliothek online verfügbar

Table of Contents

Summary ......................................................................................................... 1

Zusammenfassung ......................................................................................... 3

General Introduction ..................................................................................... 5

The peripheral nervous system ................................................................................................. 5

Injuries to the peripheral nervous system................................................................................ 6

Peripheral nerve repair ............................................................................................................. 7

Aim and structure of the thesis ................................................................................................. 8

Chapter I

Theoretical background on peripheral nerve repair and design strategies for

artificial implants ............................................................................................. 9

1 Introduction ............................................................................................................................. 9

2 Physiological requirements of artificial nerve bridges ...................................................... 10

2.1 Peripheral nerve regeneration .......................................................................................... 10

2.2 Secreted signals in peripheral nerve regeneration ........................................................... 12

2.3 Surface-bound signals ...................................................................................................... 14

2.4 Desired properties of the scaffold .................................................................................... 15

3 Design strategies for artificial nerve bridges ...................................................................... 17

3.1 Hollow tubes and multichannel nerve conduits ............................................................... 17

3.2 Structured implants with polymeric fibers ....................................................................... 21

3.3 Structured implants with gels and scaffolds .................................................................... 24

3.4 Implantation of cells with artificial nerve bridges ........................................................... 25

4 Materials for artificial nerve conduits ................................................................................. 26

4.1 Natural materials .............................................................................................................. 26

4.2 Synthetic polymers .......................................................................................................... 30

4.3 Functionalization with ECM proteins and peptides ......................................................... 35

4.4 Controlled release of growth factors ................................................................................ 37

5 Conclusion.............................................................................................................................. 39

5.1 Standardized control of efficiency ................................................................................... 39

5.2 Structured three-dimensional implants ............................................................................ 40

5.3 Biomimetic functionalization of implants ....................................................................... 40

5.4 The thirty millimeter mark ............................................................................................... 40

Chapter II

Manufacture and characterization of three-dimensional fiber arrays and

assessment of guidance properties in vitro .................................................... 41

1 Introduction ........................................................................................................................... 41

2 Material and Methods .......................................................................................................... 42

2.1 Electrospinning of three-dimensional fiber arrays .......................................................... 42

2.2 Physical characterization of fiber arrays ......................................................................... 44

2.3 Inserting the 3D fiber arrays into a conduit ..................................................................... 45

2.4 Cell and organ culture ..................................................................................................... 45

2.5 Immunocytochemistry ..................................................................................................... 46

2.6 Fluorescence microscopy ................................................................................................ 46

2.7 Quantification of Schwann cell migration and axonal growth ........................................ 47

3 Results .................................................................................................................................... 47

3.1 Electrospinning of parallel fibers in a 3D configuration ................................................. 47

3.2 Characterization of electrospun 3D PCL-fiber arrays ..................................................... 48

3.3 Suspension and manipulation of 3D fiber arrays in biocompatible gels ......................... 48

3.4 Influence of PCL fibers on axonal growth ...................................................................... 51

3.5 Influence of PCL fibers on glial cell migration ............................................................... 53

3.6 Incorporation of gel/fiber composites in a tubular conduit ............................................. 53

4 Discussion .............................................................................................................................. 54

4.1 Three-dimensional configuration of guidance structures ................................................ 55

4.2 Biological functionalization of the growth substrate....................................................... 56

5 Conclusion ............................................................................................................................. 57

Chapter III

Assembly of artificial nerve implants and assessment of their suitability to

support nerve regeneration in vivo ................................................................. 59

1 Introduction .......................................................................................................................... 59

2 Material and Methods .......................................................................................................... 60

2.1 Manufacture of artificial implants ................................................................................... 60

2.2 Animals and surgical procedure ...................................................................................... 61

2.3 Tissue preparation and immunohistochemistry ............................................................... 62

2.4 Image analysis ................................................................................................................. 63

2.5 Electromyography ........................................................................................................... 63

2.6 Behavioral experiments ................................................................................................... 64

2.7 Statistical analysis ........................................................................................................... 64

3 Results .................................................................................................................................... 65

3.1 Artificial nerve implants and general observations ......................................................... 65

3.2 Tissue regeneration: macroscopic evaluation of the implants ......................................... 66

3.3 Schwann cell infiltration and axonal regeneration in the implants ................................. 66

3.4 Size distribution of regenerated fibers ............................................................................. 70

3.5 Functional regeneration: electromyography .................................................................... 71

3.6 Functional regeneration: recovery of motor functions .................................................... 75

3.7 Muscular atrophy and regeneration ................................................................................. 77

4 Discussion .............................................................................................................................. 78

4.1 Regeneration in the artificial nerve implants ................................................................... 79

4.2 The importance of an internal guidance structure in artificial implants .......................... 80

4.3 Implementation of biological signals in synthetic cell-free nerve grafts ......................... 81

5 Conclusion.............................................................................................................................. 82

General Discussion ....................................................................................... 83

Comparison to other studies ................................................................................................... 83

Possible augmentation of implant design ............................................................................... 85

Conclusion .................................................................................................... 87

References ..................................................................................................... 89

Abbreviations ............................................................................................. 107

Figures and Tables ..................................................................................... 109

Prior Publications and Contributions ...................................................... 111

Acknowledgments ...................................................................................... 113

Curriculum Vitae ....................................................................................... 114

1

Summary

In this thesis I documented the development of an artificial nerve conduit and subsequent

biological testing within an animal model of peripheral nerve injury. The focus for the

development of the conduit lied in the utilization of electrospun microfibers to serve as a guidance

structure for Schwann cells and axons during regeneration. Through the incorporation of such a

guidance structure into the lumen of a conventional hollow nerve conduit, Schwann cell migration

and axonal elongation across artificial nerve conduits was supposed to be facilitated, thereby

expanding the applicability of artificial nerve conduits to the repair of larger nerve gaps.

To achieve this I initially established a procedure to produce three-dimensional arrays of aligned

microfibers. Through embedding of these arrays into a biocompatible hydrogel I created a

scaffold with an integrated guidance structure, which can be incorporated into a conventional

nerve conduit. This scaffold was then used to confirm in vitro, whether the guidance properties,

the fibers exerted under two-dimensional conditions, are preserved in a three dimensional

environment. Since the hydrogel (e.g. collagen) could potentially compete with the fibers for the

interaction with cells due to its fibrillary structure.

Subsequently the new implants were assembled through incorporation of the scaffold into

conventional conduits and the benefit of the integrated structure was then evaluated in comparison

to the autologous nerve graft and the conventional hollow conduit by using an animal model of

peripheral nerve injury. For this, a sciatic nerve lesion was performed in rats (Rattus norvegicus),

which was subsequently repaired by insertion of the respective im-/transplant. Assessment of

recovery by means of immunohistological investigation of the regenerated nerve,

electromyography, determination of muscle weight as well as different behavioral tests, revealed

an advantage of the internally structured conduit over the conventional hollow conduit,

particularly when the integrated fibers contained proteins derived from the extracellular matrix.

Thus, even though the autologous nerve graft still performed considerably better than the other

implants tested, the new structured design offers several possibilities for further optimization and

thus presents a promising approach to narrow the gap between autologous nerve grafts and their

artificial alternatives.

3

Zusammenfassung

In der vorliegenden Dissertation ist die Entwicklung eines künstlichen Nervenimplantats bis hin

zur Überprüfung der Wirksamkeit innerhalb eines Tierversuchs dokumentiert. Der Schwerpunkt

bei der Entwicklung des künstlichen Nervenimplantats lag in der Verwendung von Mikrofasern

als Leitstruktur für Schwannzellen und Axone während der Regeneration. Durch Integration

dieser Leitstruktur in das Lumen eines konventionellen Nervenconduits sollte die Wegfindung

besagter Zellen über das Conduit verbessert, und die Anwendung künstlicher Implantate zur

Überbrückung größerer Läsionen ermöglicht werden.

Um dies zu erreichen, wurde zunächst ein Verfahren zur Erzeugung parallel ausgerichteter

Mikrofasern in dreidimensionaler Anordnung entwickelt. Durch Einbetten der Fasern in ein

Hydrogel wurde eine Matrix mit integrierten Leitstrukturen erzeugt, die in ein konventionelles

Conduit integriert werden kann. Mit dieser Matrix wurde zunächst in einem in vitro Experiment

überprüft, ob die Fähigkeit der Fasern, die Wuchsrichtung von Zellen vorzugeben, auch in einer

dreidimensionalen Umgebung erhalten bleibt. Da diese aufgrund ihrer Struktur potentiell mit den

Fasern um die Interaktion mit den Zellen konkurrieren könnte.

Anschließend wurde die Wirksamkeit des neu entworfenen Implantats im Vergleich zum

autologen Transplantat und einem konventionellen Conduit, in einem Tierversuch überprüft.

Hierzu wurde eine Läsion des Nervus ischiadicus in der Ratte (Rattus norvegicus) gesetzt, durch

das entsprechende Im-/Transplantat überbrückt, und der regenerative Erfolg über einen Zeitraum

von zwölf Wochen mittels immunhistologischer Färbung der Nervenregenerate,

Elektromyographie, Bestimmung des Muskelgewichts der betroffenen Gliedmaßen und

verschiedener Verhaltensexperimente bewertet. Dabei zeigte sich eine Überlegenheit des von uns

entworfenen Implantats gegenüber dem konventionellen Implantat, insbesondere bei

Anreicherung der Fasern mit Proteinen der extrazellulären Matrix. Allerdings war die

Überlegenheit des autologen Transplantats gegenüber künstlichen Implantaten auch in unseren

Experimenten augenscheinlich. Nichtsdestotrotz scheint der hier vorgestellte Ansatz, bezüglich

der Struktur künstlicher Implantate, angesichts zahlreicher Möglichkeiten zur weiteren

Optimierung, geeignet den Abstand zwischen autologem Transplantat und künstlichen

Implantaten zu verringern.

The peripheral nervous system

5

General Introduction

In general, peripheral nerve injuries (PNI) are not very frequent. Different studies and statistical

surveys show an incidence of approximately two percent for injuries to the extremities (Robinson,

2004; Taylor et al., 2008) accompanied by a prevalence of around three percent (Noble et al.,

1998). Nevertheless, their socio-economic implications should not be underestimated.

First of all, there is thought to be a relatively high number of cases were the diagnosis of the

actual PNI is delayed, because shortly after an accident the treatment of the apparent injuries is

usually in the focus. Thus a considerable number of cases might fail to enter the statistics

correctly. Furthermore data from the Swiss accident analysis and statistics from 1997 (UVG)

show that the duration of working inability caused by injuries to nerves and the spinal cord is

nearly twice as long as compared to injuries in general (26,8 days vs 14,7 days). Additionally PNI

are mostly affecting the upper extremities. Thus the functional impairment in particular of the

hands can notably reduce the patient’s professional capacity for a long time and over a wide range

of working fields. This is displayed by the high follow-up costs of such injuries. For injuries to

the median and ulnar nerve, lost production makes up 87 percent of the total costs and 67 percent

for injuries to the hand and forearm in general (Rosberg et al., 2005). On top of that, the patients

suffering from PNI are relatively young with most patients between 20 and 40 years of age (Noble

et al., 1998; Eser et al., 2009; Ciaramitaro et al., 2010). For this group of people, the functional

impairment and the accompanying loss of personal independence, together with the high risk of

enduring neuropathic pain, strongly diminishes the perceived quality of life.

To explain why the treatment of these injuries presents such a challenge to the physicians in

charge, I will briefly described the causes, consequences and possible therapies of PNI in the

following.

The peripheral nervous system

The peripheral nervous system (PNS) comprises all nerves and ganglia not located within the

brain or the spinal cord, i.e. the cranial nerves (except of the optic nerve) and the spinal nerves.

The main function of the PNS is the relay of sensory information from the periphery to the central

nervous system (CNS) and the propagation of information from the CNS to the musculature and

other organs in the periphery. It can be divided into the somatic nervous system enabling

voluntary control of skeletal musculature and the autonomic nervous system mainly involuntarily

influencing internal organs and related body functions. Thus a peripheral nerve (PN) can,

depending on the innervated target organ, consist of a composition of afferent sensory axons and

efferent motor or effector axons, which according to the individual fiber type, are more or less

enwrapped by the cell membrane of Schwann cells, the myelinating glia cell of the PNS. The


6

single axons are surrounded by the endoneurium, a connective tissue forming an endoneurial tube

along the complete length of the axon. Multiple axons are ensheathed by a lamellar arrangement

of perineurial cells (the perineurium) forming a fascicle. Several of these fascicles, together with

blood vessels, are embedded in connective tissue, called epineurium, to form a complete PN.

Injuries to the peripheral nervous system

The PNS is the essential relay station for the interaction of vertebrates with their environment and

the regulation of their homeostasis. However, unlike the central parts of the nervous system, i.e.

brain and spinal cord, which are surrounded by the cranium and the vertebrate column, the PNS is

not protected by bone tissue. Thus nerves of the PNS are particularly susceptible to traumatic

injuries, especially if located in the extremities. The main cause for these injuries are traffic

accidents, followed by gunshot and cutting wounds, and are primarily affecting the upper

extremities as confirmed by an international survey in 49 countries (Scholz et al., 2009). The

consequence of a PNI is a loss of function of the innervated organ or muscle. The extent and

duration of this loss of function is essentially dependent on the position and severity of the trauma

and also on the applicable therapy.

The mildest form of PNI is a neuropraxia according to Seddon (Seddon, 1943) or a grade one

injury according to Sunderland (Sunderland, 1951). It can be caused for example by a nerve

compression that results in a damage of the myelin sheath and thus impairs signal conduction.

Since the nerve fibers as such stay intact, there usually is a rapid recovery of function after nerve

conduction is restored.

If the axon itself is damaged, but the surrounding connective tissue stays intact, one talks about

axonotmesis (Seddon, 1943) or Sunderland grade two. In this case degenerative processes will be

initiated, which for example lead to a break down and degradation of myelin and axonal parts

distal to the injury. This is followed by axonal regeneration from the proximal stump towards the

original site of innervation. The speed of functional recovery after axonotmesis is obviously

limited by the growth rate of the regenerating axons and thus does strongly depend on the distance

between site of injury and the target organ. However, as the still intact endoneurial tube serves as

a guidance structure for the regrowing axons, the prospects for functional recovery are relatively

high.

The most severe case of a PNI is called neurotmesis according to Seddon, which is further divided

by Sunderland in grades three to five and ranges from transection of the axon and its endoneurial

tube up to a complete transection of the nerve. In the milder cases of neurotmesis, where damage

to connective tissue structures is limited, regeneration is possible without surgical intervention,

but there is increasing probability of aberrant sprouting and neuroma formation. In the case of a

Peripheral nerve repair

7

complete nerve transection, however, surgical intervention is obligatory to preserve a chance for

functional recovery.

Peripheral nerve repair

A complete nerve transection results in a gap between the two nerve stumps. They have to be

reconnected by a surgical intervention to enable regeneration. If the resultant gap is small (less

than a few mm), e.g. after a clean cut, it can be sufficient to bring both stumps together and secure

them by end to end neurorrhaphy. However, frequently the amount of tissue damage or delayed

repair, which generates the necessity for trimming the nerve stumps to get a plane surface, leads

to nerve gaps considerable bigger than that. In this case continuity must not be restored by a

neurorrhaphy, since this would apply strain on the nerve, having detrimental effects for

regeneration. Thus, if relocation of the nerve to avoid strain is not possible, the insertion of a

nerve bridge is necessary (005/010 S3-Leitlinie, 2013). For this purpose, several biological

materials, e.g. muscle or vein grafts, are conceivable and their applicability is experimentally

investigated (Meek et al., 2004; Nijhuis et al., 2013). Despite this, the standard procedure is the

insertion of an autologous nerve graft, which usually is a piece of sensory nerve, most frequently

the N. suralis, N. saphenus, N. cutaneus antebrachii medialis or lateralis (005/010 S3-Leitlinie,

2013). This has, by far, the best prediction for a recovery of function, but comes along with major

drawbacks. These are, a limited availability of donor material, which more or less applies to most

biological materials, and a loss of function at the donor site, which is a particular problem of

autologous nerve grafts. To overcome these constrictions, there has been continuous research in

the past decades to test and evaluate different artificial materials in their ability to substitute the

established autologous nerve graft. It has led to the availability of clinically approved artificial

nerve conduits of different biodegradable, non-toxic materials that hardly exhibit complications,

e.g. encapsulation or increased inflammatory reactions (Meek and Coert, 2008a). However, while

these nerve conduits show satisfactory results when used to bridge small nerve gaps, the predicted

functional recovery dramatically decreases with increasing gap size, compared to the autologous

nerve graft. Therefore 30 mm are usually considered as the maximum gap size to be bridged via

tubulization with artificial nerve conduits (Meek and Coert, 2012, 2014). Generally this is

attributed to a lack of guidance cues and trophic support. An autologous nerve graft offers

thousands of endoneurial tubes that regrowing axons can follow to reach the distal stump, while

being supported by growth factors released from Schwann cells inside the graft. Compared to that

directional support in the conventional artificial nerve conduit is very poor, as they consist of

hollow tubes with a saline filling. For this reason, a main focus of current research regarding

peripheral nerve repair is the implementation of structural guidance cues into artificial nerve

conduits.


8

Aim and structure of the thesis

The goal of the project was to develop and test a cell-free artificial nerve conduit for peripheral

nerve repair that made use of electrospun, sub-micron scale fibers to improve axonal guidance. In

chapter I of the thesis I reviewed the theoretical background regarding biological processes

involved in peripheral nerve injury and regeneration as well as the current state of research in

peripheral nerve repair. Particular emphasis was put on design strategies for the improvement of

axonal guidance, the choice of materials (biological and non-biological) for artificial nerve

conduits and the incorporation of biological signaling molecules.

To achieve the experimental goal of an artificial nerve conduit the following milestones were to

be reached:

1. Designing an electrospinning device that allowed producing three-dimensional (3D)

arrays of parallel fibers for axonal guidance. This included the characterization of fiber

properties associated with their applicability for peripheral nerve repair, e.g. fiber

diameter, density and degree of alignment.

2. Manufacturing a 3D scaffold which incorporated the guidance fibers. This included,

finding a biocompatible hydrogel with suitable physical properties and establishing a

procedure for the embedding of the fiber arrays in the gel.

3. Verification of the biological properties, specifically with respect to axonal guidance and

cell migration, of the 3D fiber/gel construct. Guidance properties of the fibers used had

been shown on 2D substrates in previous studies, but needed to be confirmed for the 3D

scaffolds with appropriate cell types in vitro.

4. Integration of the 3D scaffold into an implantable tube to be tested in vivo. This device

was intended to allow for long-term storage and should be quickly available if needed.

5. Biological testing of the conduit: the efficacy to support peripheral nerve regeneration

was to be tested in vivo using the sciatic nerve lesion model in rats. Regenerative success

was to be evaluated:

a. On functional level with behavioral experiments, i.e. toe spread analysis and grip

strength test

b. On physiological level with electromygraphical recordings from the foot

musculature after stimulation of the nerve

c. By preparation of immunohistological stainings of the regenerated nerve tissue

and determination of muscle weights in the affected limb.

The experiments associated with the above mentioned milestones one to three are presented in

chapter II of this thesis, while the respective experiments for milestones four and five are

presented in chapter III.

1 Introduction

9

Chapter I

Theoretical background on peripheral nerve repair and design

strategies for artificial implants

1 Introduction

Peripheral nerves transmit motor, sensory and autonomic information between the central nervous

system (CNS) and the rest of the body. If a peripheral nerve is severed, e.g., by an injury to the

face or limbs, these functions are lost. However, in contrast to the mostly non-regenerative

processes after spinal cord lesions, axons in the peripheral nervous system (PNS) are able to

regenerate, and the long-term functional outcome of peripheral nerve injury depends on the

severity of the trauma (Sunderland, 1951). The cell somata of PNS neurons are localized in the

ventral horn of the spinal cord (motor neurons), the dorsal root ganglia (DRG; sensory neurons),

or in the sympathetic chain ganglia close to the spinal column (neurons of the autonomic nervous

system). Damage to mammalian peripheral nerve fibers activates a growth program in the

axotomized neuronal cell bodies such that their axons regenerate, provided they find a growth-

permissive substrate. The ideal growth-promoting substrate is provided by the distal segment of a

lesioned peripheral nerve.

For this reason, peripheral nerves can be surgically repaired after simple transection injuries by

reconnecting the individual proximal and distal nerve fascicles. However, when the gap between

disrupted nerve stumps is too large, the surgeon is required to transplant a segment of an

autologous nerve that is taken from elsewhere in the patient (Pabari et al., 2010). Since sensory

nerves such as the sural or peroneal nerve are often used for such procedures, the inevitable

consequence is the loss of sensory function. Although sensory nerves are routinely transplanted

even for the repair of motor fibers, experimental data indicate that sensory grafts are less suitable

than motor nerves (Brenner et al., 2006). The limitation of available donor material for

autografting and the additional risk of donor site morbidity prompted the need for alternatives to

the autologous nerve transplantation. Although several alternative tissues have been investigated

for their potential to support PNS repair, such as muscles or veins (Meek et al., 2004), a major

research focus has been the development of a bioengineering approach to design artificial nerve

conduits. Recent advances include internal topographical features that assist in the guidance of

axonal regeneration. Among these developments, the use of polymeric nanofibers has provided

one of the most promising templates for the purpose of peripheral nerve repair (Deumens et al.,

2010; Gu et al., 2011).

Chapter I

10

2 Physiological requirements of artificial nerve bridges

2.1 Peripheral nerve regeneration

To identify the requirements that artificial nerve constructs must meet in order to substitute

autologous transplants, a good understanding of peripheral nerve regeneration is important.

Peripheral nerve injury invariably causes degeneration of the distal nerve stump. Since axons are

disconnected from their cell somata they are eventually degraded. Cytoskeleton and cell

membranes are broken up into their molecular constituents. The PNS glia, the Schwann cells,

shed their myelin. This anterograde degeneration of the distal nerve segments is referred to as

Wallerian degeneration. It coincides with the infiltration of hematogenous macrophages, which

clear myelin fragments and neuronal debris. When Schwann cells lose contact with living axons

they dedifferentiate and proliferate within endoneurial tubes of the nerve. In the process, the

aligned Schwann cells form the so-called bands of Büngner, which provide an excellent growth

substrate for axonal regeneration. At the same time, retrograde signals activate a physiological

program of regeneration in the neurons. Thus, growth cones are formed by severed axons in the

proximal nerve stump (Figure 1a, b).

The success of this regenerative response depends on the extent of the injury, especially on the

maintenance of connections between the proximal fiber fascicles and the endoneurium of the

severed, distal segments. This kind of injury (called axonotmesis) can be treated conservatively

because many injured neurons survive, and regenerating axons elongate in contact with the bands

of Büngner (Sulaiman et al., 2005). Injured peripheral nerves, especially the basement

membranes, provide an excellent growth substrate for axonal regeneration, where neurite

extension reaches velocities of several millimeters per day. Subsequently, the Schwann cells re-

myelinate regenerated axons (Boyd and Gordon, 2003). If axons reach their peripheral targets

they may form new synapses or end organs, and physiological function can be restored (Figure

1c, d). When the continuity of the endoneurial sheaths is disrupted, if fiber fascicles still remain

connected via the perineurium or epineurium (neurotmesis) recovery is also possible without

surgical intervention. After complete transaction, the stumps of the elastic nerves retract, and fiber

fascicles must be re-aligned and sutured (end-to-end neurorrhaphy) (Sunderland, 1951). Larger

nerves are supported by endogenous and exogenous blood vessels, and these, too, are surgically

restored. The problems described above may represent a microsurgical challenge, but do not call

for a tissue engineering approach. This is the case, however, when the injury causes larger gaps in

the nerve, where end-to-end suturing would lead to excessive tension, impair microvascular blood

flow, and result in scarring (Pabari et al., 2010). Any larger lesions are repaired with autologous

nerve grafts (Deumens et al., 2010) because no better substrate for nerve regeneration is known

than the injured peripheral nerve itself.


11

Consequently, natural peripheral nerves provide the ideal template for biomimetic designs of

artificial nerve conduits. Ideally, such constructs would also include molecular signals that guide

and support regeneration in the natural environment. Since research on molecular processes

associated with peripheral nerve regeneration has accumulated an immense corpus of data over

the last half century, I can only mention those pertinent signal transduction pathways that have

been considered in the functionalization of nerve constructs.

Figure 1 Peripheral nerve degeneration and regeneration.

(a) Neurons of the PNS are located in the ventral horns of the spinal cord (motor neurons), the dorsal

root ganglia (sensory neurons), or the sympathetic chain ganglia (autonomic nervous system). Axons

are myelinated by Schwann cells. (b) After nerve injury myelin sheaths and axons degenerate distal

to the lesion site. Schwann cells proliferate, and macrophages remove the debris of degenerating

fibers. At the lesion site, neurons form axonal growth cones. (c) Axons are able to regenerate along

longitudinal bands of glia and ECM. Subsequently, Schwann cells remyelinate the new axons. (d) After

complete transaction proximal and distal nerve stumps retract. Frequently, when axon sprouts fail to

cross the site of injury, a neuroma forms. The distal part of the nerve and muscle fibers that are no

longer innervated become atrophic

Chapter I

12

2.2 Secreted signals in peripheral nerve regeneration

Axonal injury causes depolarization and action potentials, resulting in an influx of Ca2+

ions and

the activation of an intracellular signal transduction machinery, which regulates formation of the

growth cone (Makwana and Raivich, 2005). Once axonal growth is initiated, cell survival and

continued axonal regeneration depend on the supply of trophic physiological growth factors. Most

of these are synthesized by Schwann cells and, to a lesser extent, by macrophages and possibly

fibroblasts. Provided that the different subpopulations of peripheral neurons are maintained,

interactions with the extracellular matrix (ECM) and with other cells guide elongation of the

nerve fibers. Attempts have been made to incorporate molecular signals that mediate these

processes in tissue engineered nerve grafts (Gu et al., 2011).

2.2.1 Regeneration signals that activate axonal growth

Following the immediate consequences of injury-induced depolarization, the effect of growth

factors sets in. Perhaps the single most important signaling molecules in this category are ciliary

neurotrophic factor (CNTF) and nerve growth factor (NGF). The neurocytokine CNTF is strongly

expressed by myelinating Schwann cells. Although expression of this factor declines after the

injury, it seems to be released from the damaged cells and triggers neuronal regeneration

(Sendtner et al., 1992). Another important signal is NGF, a member of the neurotrophin family,

the expression of which is strongly reduced immediately after the lesion (Heumann et al., 1987),

and the absence of which may be important for the initiation of the regenerative response

(Shadiack et al., 2001). Similar observations were made for insulin-like growth factor-I (IGF-I)

(Kanje et al., 1991). These and related factors at later stages support survival and regeneration of

injured neurons.

2.2.2 Survival factors for the damaged nerve cells

Even in the PNS, with its high regenerative potential, a large percentage of neuronal cell death

can occur after injury, especially if the lesion occurs close to the cell soma (Deumens et al.,

2010). A number of growth factors were found to be upregulated in regenerating peripheral

nerves and, as shown by blocking their activity in vivo, to be physiologically neuroprotective.

Thus, NGF, secreted by Schwann cells, is a neurotrophic factor for sensory and autonomic

neurons. Leukemia inhibitory factor (LIF), CNTF and glial-derived neurotrophic factor (GDNF)

support motoneurons. Related molecules, especially the neurotrophins brain-derived neurotrophic

factor (BDNF), neurotrophin-3 (NT-3), NT-4/5, the neuropoietic cytokines, e.g. interleukin-6 (IL-

6), and fibroblast growth factors (FGF-1, FGF-2) have also been found to enhance survival of

PNS neurons (Heumann et al., 1987; Sendtner et al., 1990; Huang and Reichardt, 2001). Most of

these molecules have been employed in attempts to support regeneration of peripheral nerve (PN)

injury. A complementary strategy to the supply of survival factors would be to interfere with


13

apoptotic pathways that are also initiated by nerve injury. To my knowledge this has not been

tested so far in the context of repair strategies with artificial nerve constructs.

2.2.3 Promoters of axonal regeneration

Many investigations have been conducted in search of regeneration-promoting molecules in vitro

and in vivo. Important candidates that were found to be involved in the physiological context are

IGF-I and IGF-II (Glazner et al., 1993; Emel et al., 2011), both of which are also beneficial when

given exogenously. BDNF and GDNF are important growth promoting

factors for motor neurons, especially in the case of chronic deafferentiation (Sendtner et al., 1990,

1992; Boyd and Gordon, 2003). Although a sufficient supply of these factors in vivo renders

additional pharmacological application useless to the injured nerve (Gordon, 2010), they may be

important tools for artificial nerve grafts. NGF is a powerful neuritogenic factor for sympathetic

neurons (Hoyle et al., 1993). It must be noted that neurotrophins can also have neurotoxic effects

when they act via the common low affinity NGF (p75) receptor instead of the specific receptor

tyrosine kinases (trkA, trkB, trkC) (Boyd and Gordon, 2003). Receptors for several GDNF family

members (c-ret and co-receptors GFRa2, GFRa3) are expressed in small unmyelinated sensory

neurons in the DRG and in sympathetic ganglia. Their signals enhance regeneration of

sympathetic and nociceptive neurons (Trupp et al., 1995; Keast et al., 2010).

2.2.4 Modulators of the Schwann cell response

Although the effect of neurons on Schwann cells and other non-neuronal cells in the nerve are

mediated to a large extent by cell surface contacts, soluble factors play a major part. The most

important inducer of Schwann cell differentiation and myelination is neuregulin-1 of the glial

growth factor (GGF) family. GGFs are expressed by neurons and have various glia and muscle

cells as recipients. In the PNS, neuregulin-1 can drive the entire pathway of Schwann cell

differentiation including axonal myelination (Nave and Salzer, 2006). Neuregulin receptors on the

Schwann cell are ErbB2/ErbB3 heterodimers, where the ErbB3 subunit binds the ligand while

ErbB2 has a kinase activity, which initiates the intracellular signaling cascade (Citri et al., 2003).

Peripheral nerve injury induces expression of neuregulins and their receptors (Carroll et al.,

1997). Other soluble signals include nuclear receptor ligands such as thyroid hormone, estrogen,

and retinoic acid. Triiodothyronine, estrogen, and progesterone enhance axonal regeneration after

peripheral nerve injury, though it is questionable whether these hormones are endogenous signals

in the process (Panaite and Barakat-Walter, 2010). Retinoic acid signaling, which is activated by

sciatic nerve crush (Zhelyaznik et al., 2003; Zhelyaznik and Mey, 2006) promotes peripheral

nerve regeneration (Taha et al., 2004) and seems to be a crucial downstream effect of the action of

neurotrophins (Corcoran and Maden, 1999), especially for NGF- and NT-3-dependent sensory

neurons (Corcoran et al., 2000). However, retinoid receptor expression is most prominent in

Schwann cells and macrophages. The transcription factor Krox20, which is an important regulator

Chapter I

14

of peripheral myelination, is activated by neuregulin and retinoic acid via a retinoid X receptor

(Latasa et al., 2010).

2.2.5 Signals for vascularization

Members of the large, heparin-binding fibroblast growth factor family, especially FGF-1 and

FGF-2 are strong promoters of angiogenesis. Another important paracrine factor in angiogenesis

is vascular endothelial growth factor (VEGF), which acts downstream of the FGFs, but also has

parallel, independent effect (Pola et al., 2004; Murakami and Simons, 2008). In addition to these

primary regulators, synergistic interaction with other factors, including granulocyte colony

stimulating factor and platelet-derived growth factor, influence blood vessel formation in

regenerating peripheral nerves (Pan et al., 2009).

2.3 Surface-bound signals

If artificial nerve implants are to mimic the physiological stimuli that initiate and maintain

regeneration in vivo, the most promising molecular candidates would be extracted from this list of

secreted molecules for intercellular communication. In addition, since axonal growth cones

advance only in contact with surfaces, regeneration and the infiltration of growth-promoting cells

depend on the presence of permissive substrates, which interact via surface-bound signals with

cell membrane receptors. Such substrates are (1) the ECM and (2) the plasmalemma of other

cells.

2.3.1 Extracellular matrix

The predominant growth-permissive surface that renders peripheral nerves such an ideal growth

substrate is the basal lamina laid down by the Schwann cells. It consists largely of laminin and

collagen IV (Koopmans et al., 2009). Laminin and collagen are therefore the most frequently used

ECM proteins in cell culture and three-dimensional constructs (see below). Nerve fibers are

grouped in distinct fiber fascicles, which are enclosed by a perineurial matrix (endo-, peri- and

epineurium). This matrix consists of fibrillar collagens type I, III, and V, and of fibronectin,

which serves as a connecter between cell surface receptors and collagen. The cellular receptors

for ECM molecules are heterodimeric transmembrane proteins called integrins (Geiger et al.,

2001). Axonal growth cones express integrins that are activated by specific epitopes of these

ECM molecules.

Functional integrin receptors consist of one a- and one b-subunit. On one hand, they interact with

the actin cytoskeleton and, on the other hand, they can activate numerous intracellular signaling

pathways. Among the two dozen integrins known today the β1-, αv- and α4-containing integrins

are particularly important as receptors of peripheral nerve ECM. They have multiple functions,

including the regulation of cell adhesion, cell motility, and differentiation (Hynes, 1992; Geiger et

al., 2001). A synergism of neuregulin-1, expressed by axons, and the laminins in the basal lamina


15

surrounding the Schwann cells is crucial for regulation of Schwann cell differentiation and

myelination (Nave and Salzer, 2006).

2.3.2 Surface-mediated cell–cell interactions

Several classes of molecules mediate communication between axons, Schwann cells, and other

cells in the PNS. Since the late 1970s the IgG-domain containing cell adhesion molecules (CAM)

are known, especially for their role in axonal growth. Important representatives are L1/Ng-CAM,

N-CAM and transient axonal glycoprotein-1, which appear to support axon fasciculation and

axonal growth on the surface of non-neuronal cells, including Schwann cells and fibroblasts

(Bixby et al., 1988; Martini and Schachner, 1988; Martini, 1994; Soares et al., 2005). Related

myelin-associated glycoprotein and protein zero are present in myelin membranes and are

important during myelination.

Cadherins are Ca2+-binding adhesion molecules that mediate homophilic interaction between

cells. Cadherins are particularly important for selective fasciculation of regenerating axons. For

instance, E-cadherin mediates attachment of unmyelinated sensory fibers and stabilizes glial

network (Hasegawa et al., 1996), and N-cadherin mediates axon–Schwann cell interactions during

regeneration (Bixby et al., 1988; Thornton et al., 2005).

A lot of research has been done on transcriptional changes and intracellular pathways that

correlate with axonal regeneration. The intricacies of lesion-induced intracellular signaling can

only concern us here in so far as these pathways have been considered as targets for

pharmacological intervention. Suffice it to say that Ras/ERK and PI3K cascades are primarily

involved, and that Rho–type GTPases (RhoA, Rac, Cdc42) and a panel of phosphorylation-

dependent transcription factors (most importantly STAT-3, ERK), determine neuronal survival,

the regenerative response of the growth cones, and also are involved in the glial responses

(Makwana and Raivich, 2005). A considerable overlap and cross-talk between the intracellular

signaling pathways of all soluble and surface-bound signals has been observed and is far from

understood.

2.4 Desired properties of the scaffold

From the description above I can deduce some functions that artificial implants must fulfill if they

are to promote nerve regeneration in vivo. A general requirement of biocompatibility, which

applies to materials implanted anywhere in the body, demands that the implanted construct must

not induce inflammatory reactions or tumor formation and is not rejected or encapsulated by scar

tissue.

2.4.1 Mechanical Properties

In addition to biochemical signals, mechanical properties are very important because irritation of

the tissue, e.g., due to stiffness of the material, may cause inflammation or fibrosis. The basic

design of virtually all artificial nerve guides is a flexible tube, either hollow or filled with interior

Chapter I

16

structures for nerve guidance. The choice of the material for the tube, the dimensions of its wall

and lumen, as well as the filling determine its biomechanical properties. Implanted devices should

have a similar degree of flexibility as the natural peripheral nerve, while at the same time they

must be stable enough not to collapse. The Young’s modulus of various mammalian nerves has

been determined, so present day scaffolds for implantation are mechanically compatible with the

surrounding tissue (Deumens et al., 2010; Gu et al., 2011).

2.4.2 Biodegradability

It is desirable that the bridge between nerve stumps remains in the body until axonal regeneration

is completed. Studies with earlier implants, which were made of nondegradable materials such as

silicone, showed some success, but also revealed disadvantages. The long-term presence of the

material can elicit an inflammatory foreign body reaction. In several human patients,

complications with silicon tube implants required follow-up surgery to remove the implants

(Dahlin and Lundborg, 2001). Thus, recent research studies and most clinical applications

concentrate on nerve scaffolds that are gradually degraded without the release of toxic products

(Meek and Coert, 2008b; Meek and Jansen, 2009).

2.4.3 Permeability for growth factors and gas exchange

Since regeneration and physiological function of the restored nerve need the exchange of gases,

water, and biological signals such as hormones or neurotrophins the artificial implant must be

sufficiently permeable. On the other hand, osmotic influx and swelling might reduce its lumen,

thereby exerting pressure against

the regenerating fibers (de Ruiter et al., 2009). This must be avoided.

2.4.4 Guidance of axon growth and Schwann cell migration

The central function of the implant, remains, of course, the guidance of regenerating axons from

the proximal to the distal nerve segment (Dalton and Mey, 2009). Consequently, the growth cones

have to recognize physical guidance structures, which promote axonal elongation by activating

specific receptors on the cell membrane. While biochemical signaling between cells and the

scaffold surface is crucial for any interaction to take place, the physical structure of the implant

must not only allow infiltration of cells and growth cones but should guide regeneration in the

longitudinal direction toward the distal end of the implant. Polymeric nanofibers oriented in

parallel can provide ideal scaffolds in this regard. Finally, axons have to enter into the distal

fascicles of the existing nerve, which have undergone Wallerian degeneration but still lead to the

peripheral targets that are to be innervated again. Thus, an additional requirement is that the

growing axons are not trapped inside the implant.

As mentioned above, Schwann cells are indispensable not only for the process of axonal

regeneration but also for the subsequent myelination and maintenance of physiological function.

3 Design strategies for artificial nerve bridges

17

Similar molecular interactions as with neurons should therefore allow migration of Schwann cells

from the host into the nerve bridge.


With respect to the properties of the implant that will determine its success, two aspects can be

distinguished: first, the physical structure of the scaffold (see Chap.I 3), and second, the

biochemical functions of its materials (see Chap.I 4). The basic structure of all artificial nerve

bridges is a tubular graft to connect the proximal with the distal peripheral nerve stump (Figure

2).

3.1 Hollow tubes and multichannel nerve conduits

In the simplest case, a hollow cylindrical tube with a single lumen is used. The original purpose of

this design was not to replace an entire nerve segment, but only to a bridge a short gap between

the proximal and distal nerve stumps, then direct ligature of the respective fiber fascicles might

cause a strain that would interfere with the natural process of regeneration. Different techniques

have been employed for fabrication of the tube, such as melt extrusion (Chiono et al., 2009),

particle leaching (Bian et al., 2009), injection molding, and electrospinning: a variety of materials

were also tested (Meek and Coert, 2008b; Gu et al., 2011).

Since a hollow tube is the simplest device conceivable as a nerve bridge, this type of implant was

soon also tried as a long-distance connector when pieces of nerve had to be replaced. Of all

designs, the empty tube has been tested by far most often in vivo (Figure 2a and 3a, b). The tube

may consist of natural (Chap.I 4.1) or synthetic materials (Chap.I 4.2) or composites. Although

case studies with a few patients have been undertaken using more sophisticated designs (Inada et

al., 2007; Fan et al., 2008), hollow tubes constitute the only nerve guide design that has been used

in larger clinical trials so far (Meek and Coert, 2008b; Gu et al., 2011). The many cell culture and

animal experiments that led to the clinical studies will not be reviewed here.

3.1.1 Empty nerve conduits

The non-degradable tubes that were first implanted in patients, consisted of silicon (Lundborg et

al., 1997b, 2004) or polytetrafluoroethylene (Stanec and Stanec, 1998; Pitta et al., 2001). In one

large study, 26 patients had suffered injuries to the median or the ulnar or to both nerves and

received tubular silicone implants as nerve bridges. Based on sensory and motor function and

pain, in 19 cases the outcome of the operation was rated very good or good; however, the implant

Chapter I

18

had to be removed in seven cases because of discomfort caused by the silicone tube (Braga-Silva,

1999). There are divided opinions about whether non-degradable implants are advisable because

they may cause a foreign body reaction, and whether the need to remove them in a subsequent

operation after regeneration is risky (Lundborg et al., 1997b). However, even Lundborg and

colleagues who have conducted successful clinical trials with silicon tubes state a preference for

Figure 2 Design strategies for artificial nerve bridges.

(a) Empty tube used for bridging gaps between proximal and distal nerve stumps; (b) tubular

implant with multiple intraluminal channels (e.g., de Ruiter et al., 2008; Yao et al., 2010); (c) aligned

microfibers as longitudinal guidance structures (e.g., Kim et al., 2008; Ribeiro-Resende et al., 2009);

(d) hydrogel filling of the conduit (e.g., Inada et al., 2004; Labrador et al., 1998); (e) sponge filled tube

(e.g., Inada et al., 2004; Toba et al., 2001); (f) longitudinally oriented pores within the scaffold (e.g.,

Ceballos et al., 1999; Möllers et al., 2009); (g) modification of the inner side of the tube with bioactive

coating (e.g., Chew et al., 2007; Shen et al., 2010)


19

biodegradable materials, provided their degradation does not cause further complications (Dahlin

and Lundborg, 2001). Thus, to date, biodegradable materials are the preferred solution. Three

types of such tubes have been approved for implantation in humans and are commercially

available. They have been tested in several hundreds of patients.

NeuraGen, consisting of type I collagen, is produced by Integra Neuroscience (www.integra-

ls.com). In the largest study so far, NeuraGen implants were used mostly for bridging sensory

nerves of the arm. Twenty-six patients were evaluated quantitatively to assess functional recovery

of nerve transmission. In 45 % of the patients an improvement of sensory functions was seen

(Wangensteen and Kalliainen, 2010). In a number of small clinical trials and case studies, the

collagen implant has demonstrated its usefulness (Ashley et al., 2006; Lohmeyer et al., 2007;

Farole and Jamal, 2008; Taras and Jacoby, 2008), though occasional failures were also reported

(Moore et al., 2009). In this latter report, another nerve conduit, Neurotube, was also tested,

though without better outcome.

Neurotube is a poly(glycolic acid) (PGA) implant, produced by Synovis Life Technologies

(www.synovismicro.com). It has been tested in several clinical studies. In one of these, a group of

46 patients received Neurotube implants for lesions of nerves innervating the hand. The results

showed good to excellent outcome in 74 % of the cases, which was not statistically different from

a control group who received autologous nerve transplantations (Weber et al., 2000). Other

successful studies with the PGA tubes have been published, with repair of median, ulnar, and

facial nerves being reported (Navissano et al., 2005; Donoghoe et al., 2007; Agnew and

Dumanian, 2010).

The third commercially available nerve scaffold is Neurolac, produced by Polyganics

(www.polyganics.com). It consists of a poly(lactic acid)/poly(ε-caprolactone) blend (PLA/PCL).

So far, two clinical studies and case reports have been published, with a total of 36 patients.

Although in the first report from 2003 no nerve regeneration was found, the second study reported

functional regeneration similar to results with autologous nerve transplants (Bertleff et al., 2005).

Again, failures to achieve recovery of sensory functions were reported, e.g., with digital nerve

reconstruction in the foot (Meek et al., 2006).

3.1.2 Multichannel devices

Without changing the basic design, tubular implants have been constructed to incorporate several

intraluminal channels that can accommodate different fiber fascicles within a peripheral nerve

(Figure 2b and 3c, d). Multichannel implants with up to seven compartments were produced using

a similar injection-molding technique as for the implant with a single lumen (de Ruiter et al.,

2008; Yao et al., 2010). On the level of fiber fascicles, the multichannel devices appeared to

reduce mixing of regenerating axons. Necessarily, these constructs reduced the available space

within the implant; and when numbers of regenerated fibers and behavioral improvement were

Chapter I

20

evaluated, the multichannel constructs were no better than hollow tubes (Yao et al., 2010). An

alternative construction design achieved the similar effect of longitudinal divisions of the

otherwise empty tube by integrating one or a few films into the construct. When tested in rat tibial

nerves, tubes with one film, i.e., just two compartments, were better than implants with none or

three films (Clements et al., 2009).

In animal experiments, many studies were performed with hollow tubular implants to assess the

biocompatibility of the material. Most often, grafts were tested as bridges of the sciatic nerve, and

by far most experiments were done with rats where nerve lesions rarely exceeded 20 mm. When

short nerve gaps are bridged with tubular implants, a fibrin matrix will fill this distance and allow

infiltration of Schwann cells from the nerve stumps. These cells can then form orientated bands of

Büngner, which may guide regenerating axons. Experiments with rabbits, cats, dogs, and

monkeys indicate that internal guidance structures are needed for larger cell-free implants (Brown

et al., 1996; Matsumoto et al., 2000; Ding et al., 2010). An extensive list of animal studies with

further references is given in a recent review (Shin et al., 2012). The construction of three-

dimensional guidance structures remains a technical challenge, where polymeric nano- and

microfibers take center stage.

Figure 3 Examples of artificial nerve guides that were successfully tested in animal

experiments.

(a) SEM images of a nerve tube consisting of electrospun PLGA/PCL microfibers; scale bar 500 µm

(Panseri et al., 2008). (b–d) Collagen conduits with one or multiple interior channels were fabricated

from molds with insertion of stainless wires; scale bar 1 mm (Yao et al., 2010). (e) Composite nerve

conduit of several hundred PCL filaments each having six leaflets in cross-section (insert); scale bar

150 µm (Whitlock et al., 2009). (f) Poly(glycolic acid) tubes filled with collagen sponge were

successfully used to bridge 80 mm gaps of canine peroneal nerves (Nakamura et al., 2004). (g)

Collagen scaffold produced with directional freezing; longitudinal sections demonstrate oriented

channels; extensive fenestration between adjacent channels can also be seen; scale bars 50 µm

(Möllers et al., 2009).


21

3.2 Structured implants with polymeric fibers

One successful strategy to provide such guidance structures in the longitudinal direction of the

implant is the use of polymer fibers (Figure 2c and 3e). These can be produced by a number of

different techniques such as drawing, template synthesis, phase separation, self-assembly, and

electrospinning. Historically, large diameter fibers were tested first. Lundberg and coworkers

filled silicone tubes with eight longitudinally oriented polyamide filaments of 250 µm diameter as

implants. Across a 15 mm gap in the rat sciatic nerve, this device was a substantial improvement

compared with hollow implants (Lundborg et al., 1997a). To mimic the natural situation, many

more and thinner fibers would provide topological cues and induce the formation of many

longitudinal strands of Schwann cells. Cell culture experiments demonstrated that small caliber

fibers (5–30 µm) had a stronger effect on the orientation of growing neurites than thicker fibers

(500 µm) (Smeal et al., 2005). In the context of peripheral nerve regeneration, the technique that

has been used most frequently for the manufacturing of small diameter fibers is electrospinning.

3.2.1 Electrospinning of nano-/microfibers

As a process to produce continuous fibers with very small diameters, electrospinning has been

known for more than a century. During the last two decades, the method has become increasingly

popular in the field of bioengineering because the method is easy to use, very flexible with respect

to fiber configuration in situ, and a great range of different materials can be used (Teo and

Ramakrishna, 2006). Biodegradable polymers that have been spun to nanofibers, tested in vitro,

and implanted in the nervous system will be discussed below. The typical electrospinning set-up

in a research laboratory consists of a high-voltage power supply (up to 30 kV), the spinneret (a

syringe with a flat tip needle), and collector electrodes (Figure 4). To induce formation of fibers, a

high voltage is applied via the spinneret to the polymer solution. This causes electrostatic

repulsion within the charged solution, which is stronger than its surface tension, such that a jet

erupts from the spinneret. At some distance, the jet enters a stage of bending instability, is

stretched under electrostatic forces toward the grounded target electrode, and the solvent

evaporates. With a low flow rate, polymer solution is pumped through the spinneret to ensure

continuous formation of fibers, which accumulate on the target. On a single plate electrode, used

as a target, mats of randomly coiled fibers accumulate. Several electrospinning devices were

tested in order to assemble nanofibers that are aligned in a parallel orientation. Rotating drums in

different configurations were devised by several groups (Matthews et al., 2002; Chew et al., 2005;

Kim et al., 2008). Alternatively, two parallel bars can be used as target electrodes, such that

parallel fibers accumulate between them (Schnell et al., 2007; Klinkhammer et al., 2009). Using

such devices, parallel fibers of diameter from less than 100 nm to more than 5 µm can be

produced from various synthetic polymers. A good review with drawings of various devices has

been written by Teo and Ramakrishna (Teo and Ramakrishna, 2006).

Chapter I

22

Electrospun micro- and nanofibers have high surface-to-volume ratios and provide growth

substrates for many neural cell types (Gerardo-nava et al., 2009; Bockelmann et al., 2011). Films

of electrospun fibers have been used to produce tubular implants and to subdivide nerve conduits

into longitudinal compartments (Clements et al., 2009). The specific advantage of nanofibers is,

of course, their property as individual guidance structures (Lietz et al., 2006; Kim et al., 2008). In

one study, PGA microfibers were incorporated in a chitosan tube as implants to bridge 30 mm

gaps in the sciatic nerve of beagles. The dogs, which were investigated after 6 months, recovered

function of the operated nerves. Skeletal muscles were re-innervated, and the artificial constructs

were completely degraded within the half year time-frame of the study (Wang et al., 2005). A

different research team stacked films of parallel fibers consisting of poly(acrylonitrile-co-methyl

acrylate) (PAN-MA) inside a polysulfone tube. These constructs were implanted in rats to bridge

sciatic nerve gaps of 17 mm. Sixteen weeks after surgery, target muscles were found to be

innervated again, and regeneration of sensory and motor nerve fibers could be demonstrated (Kim

et al., 2008). This represented a significant progress because the scientists incorporated oriented

nanofibers in a three-dimensional design. The largest distance of nerve repair has been reported

by Matsumoto, Toba and coworkers, who were able to bridge an 80 mm gap of the peroneal nerve

in dogs (Matsumoto et al., 2000). Their conduits consisted of tubes of PGA/collagen blend, filled

with laminin-coated collagen fibers. Regeneration was assessed histologically,

electrophysiologically, and with behavioral testing. Very sophisticated guidance structures have

Figure 4 Electrospinning of aligned fibers.

The electrospinning set-up used by the authors consists of a high voltage power supply (e.g., 20 kV),

the spinneret (a syringe with a flat tip needle) connected to a pump (flow rate e.g., 1 mL/h) and

parallel bars as target electrodes to collect parallel fibers (Schnell et al., 2007; Klinkhammer et al.,

2009). Another frequently used device to assemble aligned fibers is the rotating drum (Matthews et

al., 2002; Chew et al., 2005; Kim et al., 2008). Scanning electron microscopy images show

PCL/collagen fibers collected between parallel bars (Gerardo-nava et al., 2009)


23

been developed by Schlosshauer’s group, who also tested functionalization with NGF,

transforming growth factor-b1 (TGFb) and laminin in their experiments. Oriented PCL filaments

were made by melt extrusion using nozzles with six-leaf cross-sections. This resulted in yarns of

several hundred filaments, each with six longitudinal grooves (Figure 3e), which caused the

alignment of Schwann cells, “artificial bands of Büngner” (Ribeiro-Resende et al., 2009). As also

observed by others (Schnell et al., 2007), the topographic effect of the substrate on glia cells

indirectly affected the orientation of growing axons that need not have any contact with the fibers

themselves.

3.2.2 Self-assembling peptide scaffolds

The manufacture of self-assembling nanofibrous scaffolds is based on repetitive peptides that

consist of alternating hydrophilic and hydrophobic amino acids and spontaneously form stable b-

sheet structures. The self-assembling process takes place when the aqueous peptide solution is

introduced to a physiological salt-containing solution (i.e., saline, tissue culture media,

physiological solutions, or body fluids such as cerebrospinal fluid) and is mediated by non-

covalent bonds, such as van der Waals forces, hydrogen bonds, and electrostatic forces. As a

result of hydrophobic interactions, different b-sheets pack together and form double layered β-

sheet nanofibers with diameters of about 10 nm, much smaller than typical electrospun nanofibers

(Gelain et al., 2007; Cao et al., 2009). Scaffolds based on self-assembled peptides provide several

advantages over other biomaterials for their potential use as bridging materials for the nervous

system:

1. Their three-dimensional environment has dimensions that are similar to the

native ECM in peripheral nerves.

2. The peptides can be degraded into natural L-amino acids that are nontoxic and

may be recycled by surrounding cells.

3. They elicit only a minor, if any, inflammatory or immune response.

4. Modifications at the single amino acid level are possible, and various functional

motifs can be added to promote neurite outgrowth.

5. The self-assembly process takes place under physiological conditions without

the need of temperature changes (Gelain et al., 2007, 2010).

Several cell culture studies demonstrated that scaffolds made via self-assembling peptides provide

permissive substrates for cell attachment and growth as well as support for extensive neurite

outgrowth and synapse formation (Holmes et al., 2000; Silva et al., 2004; Ellis-Behnke et al.,

2006), including the migration of Schwann cells and axonal regeneration from DRG neurites

(DRG) (McGrath et al., 2010). Fewer reports have been published where such scaffolds were used

in the nervous system in vivo. In one study of CNS repair, a knife wound in the visual system of

Chapter I

24

2-day-old hamsters was successfully treated by injection of peptide solution into the lesion site.

Here self-assembling nanofiber scaffolds provided a permissive substrate for axonal regrowth,

and visual function could be restored (Ellis-Behnke et al., 2006). Peptide scaffolds implanted after

spinal cord dorsal column transection injuries of rats integrated very well with the host tissue and

supported the infiltration by blood vessels and axons (Guo et al., 2007). To my knowledge, only

one in vivo study focused on the potential of self-assembly nanofibers to repair a peripheral nerve

(McGrath et al., 2010). In this report, a 10 mm gap of the rat sciatic nerve was successfully

bridged with a tubular conduit filled with a self-assembling peptide nanofiber scaffold. The

distance of axonal regrowth 3 weeks after implantation was significantly enhanced by these

conduits when compared to control implants filled with alginate/fibronectin hydrogels, and

recovery of gastrocnemius muscle function was also better. However, grafted Schwann cells

improved the efficacy of the implants, and autologous nerve grafts still yielded the best outcome

(McGrath et al., 2010). Although these investigations demonstrated that self-assembling peptide

scaffolds can serve as permissive substrate in regenerative medicine, the nanofibers of these

constructs lack orientation. This is a desired property for long distance guidance within an

artificial nerve bridge, which is more easily achieved with the electrospinning method.

3.3 Structured implants with gels and scaffolds

A completely different approach for the internal structure of implants consists in gel scaffolds

with a narrow mesh of longitudinal pores or channels. Several laboratories used collagen

hydrogels to fill the lumen of different nerve guides, which themselves were made from a range

of biocompatible materials (Figure 2d). Based on histology and measurements of functional

recovery, such hydrogel fillings of collagen or laminin proved superior in comparison with saline-

filled implants. It is interesting to note that lower concentrations of collagen- or laminin-

containing gels (e.g., 1.28 mg/ml collagen, 4 mg/ml laminin) supported better functional recovery

than high gel concentrations (e.g., 1.92 mg/mL collagen and 12 mg/ml laminin) (Labrador et al.,

1998; Verdú et al., 2002).

Scientists from the University of Kyoto constructed a sophisticated nerve conduit that consists of

a polyglycolic acid tube, filled with a sponge of porcine collagen soaked with laminin (Figure 2e

and 3f) (Toba et al., 2001; Inada et al., 2004). Tubes of 3–4 mm inner diameter were implanted as

bridges into the peroneal nerves of beagle dogs. On the basis of histological and

electrophysiological evaluation, these implants were even superior to nerve autografts, although

some cases of neuromas occurred (Inada et al., 2004). Consequently, similar tubes were implanted

for repair in human patients. In two women, the frontal branch of the facial nerve was restored

with the construct. Five months after surgery both were able to lift their eyebrows symmetrically.

Electrophysiological tests showed recovery of compound muscle action potentials with normal

latency (Inada et al., 2007). This success was possible with the repair of small caliber nerves,


25

which also showed the best prognosis using hollow tubular implants. Nevertheless, the outcome

and the previous animal experiments, where bridges of up to 80 mm were implanted in dogs

(Toba et al., 2001; Siemionow et al., 2010), demonstrated that the inclusion of a collagen matrix

was a substantial advancement.

While these results are already very impressive, additional improvement is expected with the

design of gels with a fibrillar network that is oriented in the longitudinal direction of the nerve

guide (Figure 2f). One idea to achieve this is the application of strong magnetic fields that can

orient collagen or fibrin matrices. Such scaffolds promoted neurite elongation from chick DRG in

the desired direction (Dubey et al., 1999). In mice sciatic nerves, magnetically aligned collagen

gels supported regeneration and remyelination of axons over a 6 mm gaps better than control

tubes (Ceballos et al., 1999).

Another, very promising technique is controlled freeze drying of collagen suspensions (Figure

3g). When solvent is caused to freeze from one side of the solutions, collagen is displaced to the

side, thus resulting in walls of orientated guidance channels. The diameter of these longitudinal

pores can be controlled in the range of 20–120 µm. In vitro studies have already demonstrated

that such scaffolds can be easily infiltrated by Schwann cells, fibroblasts and regenerating axons,

which follow the desired orientation of the scaffold (Bozkurt et al., 2007, 2009; Möllers et al.,

2009). Three-dimensional constructs produced with this method are commercially available

(Matricel www.matricel.net) and have been tested successfully in animal experiments (Bozkurt et

al., 2012).

3.4 Implantation of cells with artificial nerve bridges

The closest thing to real nerve transplants are artificial constructs that are preseeded with growth

promoting cells, and many researcher believe that gaps longer than 30 mm will not be

successfully repaired unless supporting cells are included in the scaffold (Hood et al., 2009;

Deumens et al., 2010). These cells may derive from the host organism itself or from cultures.

They are intended to release physiological signals, many of which may be unknown and thus

cannot be mimicked with chemical modification of the implant material. Best suited for this

purpose are Schwann cells because they naturally populate the PNS, where they serve a number

of important functions during regeneration. As discussed above, Schwann cells secrete the basal

lamina that serves as axon guidance substrate, synthesize a host of necessary growth factors, and

even participate in the clearance of myelin debris. Later, they are required for myelination of the

regenerated fibers (Boyd and Gordon, 2003). Unfortunately, isolation of Schwann cells from a

peripheral nerve and expansion in culture are time-consuming and also carry the risk that culture

conditions alter the cellular phenotype.

Other cell sources have therefore been explored as alternatives. They include olfactory

ensheathing cells, embryonic stem cells, mesenchymal stem cells (Gu et al., 2011; Radtke et al.,

Chapter I

26

2011), and even cell lines such as the immortalized SCTM41 Schwann cells (Lago et al., 2009).

Genetic modification has been considered for implanting cells as a constant source of growth

factors. Most often, fibroblasts have been engineered with this intention, e.g., to produce

neurotrophins, CNTF and FGF-2 (Sørensen et al., 1998; Schmidt and Leach, 2003).

The majority of studies on PN regeneration, however, used Schwann cell primary cultures

(Guénard et al., 1992; Kim et al., 1994; Brown et al., 1996; Sørensen et al., 1998; May et al.,

2004; Radtke et al., 2005; Lohmeyer et al., 2007; Sinis et al., 2007). Unfortunately, the

implantation of any type of living cells is not without complications. First, the purity of the

primary culture is essential because the cells must not be contaminated, e.g., by fibroblasts in a

Schwann cell culture. Second, immune rejection of the cells is an issue. Heterologous

transplantation poses less of a problem between inbred strains of mice and rats, but it would

require immune suppression in humans. Ideally therefore, the cells would have to be taken from

the same patient who requires the PN implant. For a potential therapy of CNS lesions, an

interesting strategy is the preparation of olfactory ensheathing cells from the human olfactory

epithelium. In many physiological respects these glial cells are similar to Schwann cells (Kawaja

et al., 2009), and the cells could be prepared for autologous implantation. Olfactory ensheathing

cells are therefore being explored as a possibility to produce better implants also for peripheral

nerves (Radtke et al., 2011). Nonetheless, because of the fundamental difficulties of immune

rejection when foreign tissues are implanted, I consider the development of cell free artificial

nerve bridges as an important goal. Ideally, scaffolds of different lengths and diameters could be

produced in advance and stored to be implanted on demand.

4 Materials for artificial nerve conduits

Biocompatible materials are needed for the two essential components of the artificial nerve

bridge, i.e., the surrounding tube and its interior structure, the latter being either linear filaments

or a scaffold with longitudinal pores. The selected material will determine not only the

biochemical interaction between cells and graft but also the physical properties of the implant.

4.1 Natural materials

Many natural materials are hydrophilic and show excellent biocompatibility. Proteins and

carbohydrates of high molecular weight are extracted rather than synthesized chemically. On the

down side, when the material is derived from mammalian tissue there is a certain risk of disease

transmission and/or allergic reactions. Nevertheless, many biologically derived materials have just

the desired properties for applications in nerve repair.


27

4.1.1 Extracellular matrix proteins

The most straight forward approach is to use ECM material that gives structure to the peripheral

nerves in situ. Epineurium, perineurium and endoneurium consist largely of one family of ECM

proteins: collagen.

Collagen

Of all biologically derived material used in bioengineering, collagen is the most popular, not only

for PN conduits but for tissue reconstruction in many other applications. Collagens, which

account for more than a quarter of total body protein, are extracellular proteins with a

characteristic triple helix structure of three either identical or heteromeric a-chains (Koopmans et

al., 2009). Of at least 29 different forms, collagen I, III, and V form fibrils in the ECM of the

PNS. The non-fibrillar collagen IV is present in the basement membranes secreted by Schwann

cells. In vitro, collagen IV has superior qualities as a substrate for axonal growth, but the fibrillar

collagen I is more often used in scaffold production. As reviewed above, collagen has been used

to form tubular implants, e.g., NeuraGen (Ashley et al., 2006; Farole and Jamal, 2008; Taras and

Jacoby, 2008; Wangensteen and Kalliainen, 2010). It is also one of the materials frequently used

for interior guidance structures (Figure 3b–d, f, g): During matrix formation, collagen hydrogels

can be oriented with strong magnetic fields (Ceballos et al., 1999), and collagen sponges

combined with laminin (Toba et al., 2001; Inada et al., 2004) or chondroitin-6-sulfate

(Chamberlain et al., 1998) were very successful as nerve guides in animal experiments. Using the

directional freezing technique, scaffolds with longitudinally oriented pores were created and

tested for implantation in PNS and CNS (Bozkurt et al., 2007, 2009; Möllers et al., 2009). It is

also possible to produce collagen nanofibers by electrospinning (Matthews et al., 2002). Even

blends with other polymers (e.g., 25 % collagen, 75 % PCL) were significantly better for axon

guidance and Schwann cell migration than control fibers without collagen (Schnell et al., 2007).

Fibronectin

Axons and glia cells interact with the ECM via their integrin receptors. Although integrins can

directly bind to some epitopes on collagen, an important intermediate molecule between collagen

and the cells is the 440 kDa glycoprotein fibronectin, which is also part of the basement

membrane. It is primarily produced by fibroblasts. In addition, there is a soluble dimeric form of

fibronectin in the plasma, which together with fibrin has an important function in wound healing

(Pankov and Yamada, 2002). The biological activity of fibronectin has been shown to be strongly

influenced by the conformation of underlying surfaces (Dalton and Mey, 2009). Fibronectin

conduits have performed well in PN repair. For instance, tubular implants filled with a solution of

fibronectin, laminin, and fibronectin/laminin mixture permitted regeneration across an 18 mm gap

in rat sciatic nerves, and all of these ECM molecules promoted migration of Schwann cells into

the conduit (Bailey et al., 1993).

Chapter I

28

Laminin

This is the most important component of the basal lamina. Different laminins exist, all of which

are large glycoproteins (500–1,000 kDa) that consist of three polypeptide chains (α, β, and γ).

Applied in gels or solution together with collagen and other proteins, laminin enhanced nerve

regeneration in vivo and in vitro. In addition to the whole protein, integrin-activating peptide

sequences from laminin were identified that also mediate cell adhesion and growth. Well-

characterized sequences are those containing Arg-Gly-Asp (RGD, in one letter code) in laminin a-

chains (Hersel et al., 2003); YIGSR, located on the β1-chain; and IKVAV, which is present near

the C-terminal end of the α1-chain, close to a heparin-binding domain (Dalton and Mey, 2009).

When synthetic polymer fibers were functionalized with GRGDS, adhesion and speed of axonal

growth cones and of migrating cells were significantly enhanced (Bockelmann et al., 2011).

Laminins have already been used for a long time as substrates in cell culture experiments. The

commercially available Matrigel is a non-defined ECM solution derived from a mouse sarcoma

cell line. Its main component is laminin, but Matrigel also contains collagen IV, heparan sulfate

proteoglycans, and several growth factors. Used as a thermoreversible gel it is liquid at low

temperatures and solidifies at body temperature. As a filling in fiber containing implants, Matrigel

supported regeneration in the mouse sciatic nerve. Electromagnetic alignment of the material

further improved the outcome (Verdú et al., 2002; Lohmeyer et al., 2007).

4.1.2 Other natural proteins

Fibroin (Silk)

Silk is not immunogenic and has high tensile strength, toughness, and elasticity (Deumens et al.,

2010). Its core structural protein is fibroin, extracted from the cocoons of Bombyx mori

caterpillars (silkworm). The protein exhibited good biocompatibility with Schwann cells and

DRG (Yang et al., 2007). Fibroin was used for electrospinning of nanofibers (Madduri et al.,

2010b). A tubular fibroin construct, filled with about 20 longitudinally aligned fibroin fibers was

implanted as a 10 mm nerve bridge in rats and evaluated after 6 months. The detailed anatomical

evaluation demonstrated excellent results, though autologous nerve grafts were still superior

(Yang et al., 2007).

4.1.3 Polysaccharides

One risk with the implantation of foreign ECM proteins is development of scar tissue within the

regenerating area. This acts as a physical barrier to axonal elongation between the nerve stumps.

It appears that this problem is less likely with biocompatible polysaccharides.

Chitin and chitosan

The exoskeleton of arthropods consists largely of chitin, which is also present in the cell walls of

fungi. Second only to cellulose, chitin is the most abundant polysaccharide found in nature. The


29

monomeric building blocks of chitin are N-acetyl-D-glucosamine. Chitosan, a linear copolymer of

N-glucosamine and N-acetyl-D-glucosamine, is prepared via partial deacetylation of chitin. Due

to similarities with glucosaminoglycans side chains of ECM glycoproteins (e.g., chondroitin

sulfate, heparin sulfate, keratan sulfate), chitosan can interact with fibronectin, laminin, and

collagen. It has been used extensively in biomedical applications including nerve conduits (Wang

et al., 2005; Gu et al., 2011). In one construct, a double-layered tube was devised comprising an

outer chitosan film for mechanical strength and an inner layer of electrospun chitosan fibers

(Figure 2g). In addition, extended YIGSR peptides were covalently bound to the chitosan fibers to

enhance interaction with cells. As 10 mm peripheral nerve guides in rats, these constructs were

successful as autografts (Itoh et al., 2005; Wang et al., 2008). When chitosan/PLGA implants

were complemented with bone marrow-derived stem cells, even 50 mm sciatic gaps in dogs could

be repaired almost as well as with autologous nerve transplants (Ding et al., 2010). Chitosan

scaffolds were also prepared with longitudinal pores, which improved regeneration and functional

recovery (Huang et al., 2010). In other experiments, chitin/chitosan implants caused infiltration of

many inflammatory macrophages, and modification of the material was therefore considered

(Huang et al., 2010).

Alginate and agarose

Alginate, extracted from brown seaweed, consists of repeat units of mannuronic acid and

glucuronic acid. Having free carboxyl groups it can be cross-linked in the presence of calcium.

Most often, alginate is used as a gel inside implanted nerve tubes (Suzuki et al., 1999; Ohta et al.,

2004); however, alginate gel has been used as glue to connect nerve stumps 7 mm apart (Suzuki et

al., 2000), and hollow tubes were made of freeze-dried alginate/chitosan blend. For the delivery of

a growth factor, nerve tubes were filled with alginate and implanted in a rat sciatic nerve model.

This study showed that regeneration was better in empty conduits, suggesting that the presence of

alginate actually impeded regeneration (Mohanna et al., 2003). Agarose is also extracted from

seaweed and has been used extensively for cell cultures (Dalton and Mey, 2009). In the context of

artificial nerve implants, an interesting application has been the use of agarose as a material for

creating concentration gradients of growth factors (Mohanna et al., 2003, 2005).

Hyaluronic acid

Another important glucosaminoglycan of the ECM in connective tissues, especially cartilage and

skin, is hyaluronic acid. It consists of repeating disaccharide units of D-glucoronic acid and D-N-

acetyl-glucosamine. For biomedical applications, hyaluronic acid gels proved to be useful because

the molecule is involved in many aspects of wound repair, including angiogenesis, reduction of

scar formation, and moderation of the inflammatory reaction. A positive effect on PN

regeneration was achieved with local application (Ozgenel, 2003). Hydrogels consisting of

Chapter I

30

hyaluronic acid and ECM proteins were used as carrier matrix in cell-based repair strategies for

the PNS (Suri and Schmidt, 2010).

4.1.4 Self-assembling peptide scaffolds

As a new strategy in the design of scaffolds for nerve repair, self-assembling peptide nanofibers

have already been mentioned above. They consist of natural amino acids and form a three-

dimensional network of nanofibers through spontaneous self-assembly. The ability of EAK16-II

(AEAEAKAKAEAEAKAK; one-letter amino acid code for Glu, Ala, Lys), a 16 amino acid

peptide derived from the yeast protein zuotin, to self-assemble into a nanofibrous scaffold was

observed serendipitously when the peptide was tested for its cytotoxicity in cell culture (Zhang et

al., 1993). This finding has inspired the development of a number of additional self-assembling

peptides including RADA16-I (AcN-RADARADARADARADA-CNH2) and RADA16-II (AcN-

RARADADARARADADA-CNH2), in which Arg (R) and Asp (D) residues substitute Lys and

Glu (Holmes et al., 2000). Studies, which demonstrated the compatibility of self-assembling

peptides with neural tissue were cited above. In a few experiments, these were tested for nerve

repair (Rafiuddin Ahmed and Jayakumar, 2003; Silva et al., 2004; Ellis-Behnke et al., 2006;

McGrath et al., 2010). The RADA16 self-assembling peptide solution has been developed into a

commercial product intended for wound repair and three-dimensional cell cultures (3DM,

www.PuraMatrix.com).

4.2 Synthetic polymers

In comparison with biologically derived materials, synthetic polymers have several advantages

and disadvantages. Since they do not need to be extracted from animal tissue they pose a lower

risk of contamination and immunological reaction. Their physical and chemical properties can be

better controlled and manipulated. On the other hand, synthetic materials usually do not activate

specific cellular receptors and therefore do not by themselves elicit the desired cellular responses

such as axonal growth or cell migration. Some materials may even be rejected or encapsulated in

scar tissue.

4.2.1 Non-degradable materials

The first generation of tubular implants for PN repair were made of silicone rubber, which

showed some success (Lundborg et al., 1982, 1997b, 2004; Braga-Silva, 1999). Another non-

degradable material is poly(acrylonitrile-co-methylacrylate) (PAN-MA). Recently, a peripheral

nerve conduit was designed by stacking oriented, electrospun PAN-MA fibers in the two halves

of a longitudinally split polysulfone tube, which were resealed and implanted as a sciatic nerve

bridge in rats. In the absence of additional cells or growth factors this construct supported

excellent regeneration, which was monitored with histological, physiological, and behavioral

methods (Kim et al., 2008). This and other studies (Clements et al., 2009) demonstrated the

advantage of three-dimensional arrays of nanofibers as guidance structures in vivo (Figure 5).


31

Other non-degradable materials used for nerve implants are extended poly(tetrafluoroethylene)

(ePTFE, Gore-Tex) and poly(urethane) (Stanec and Stanec, 1998). Since non-degradable

materials may cause a chronic foreign body reaction and therefore need to be removed in a second

operation, they are no longer a preferred strategy for artificial nerve implants (Gu et al., 2011).

4.2.2 Aliphatic polyesters

Perhaps the most important category of biodegradable synthetic polymers with favorable

characteristics for nerve repair is that of aliphatic polyesters. The most important examples are

poly(a-hydroxy acids), i.e., poly(lactic acid) (PLA, with stereoisomers PLLA, PDLA),

poly(glycolic acid) (PGA), poly(lactic-co-glycolic acid) (PLGA), poly(ε-caprolactone) (PCL), and

poly(hydroxybutyric acid) (PHB). The compounds are completely biodegradable via hydrolytic

digestion, and many cell culture experiments demonstrated excellent biocompatibility with nerve

tissue of PLA (Yang et al., 2005; Liu et al., 2011), PGA (Ichihara et al., 2009), PLGA (Bini et al.,

2004; Ding et al., 2010; Yucel et al., 2010), PCL (Schnell et al., 2007; Klinkhammer et al., 2009;

Wang et al., 2009), and PHB (Mohanna et al., 2003, 2005; Aberg et al., 2009). However, they are

not soluble in water, do not directly activate biological signals, so possibilities for conjugation

with functional peptides are limited. For nerve bioengineering purposes, poly(a-hydroxy acids)

are dissolved in non-aqueous solvents and spun to fibers that can be used as material for the nerve

tube or as oriented guidance structures for growing axons and cells. Fibers can also be spun

directly from the polymer melts. Often blends of different polymers or blends with biologically

derived materials have been used successfully. A few important examples are given below, but, as

many other studies could be cited, this selection is somewhat arbitrary.

Poly(L-lactic acid)

Cai and coworkers produced porous tubular PLA conduits, which had an internal structure of

aligned PLA filaments. Ten weeks after implantation in a rat sciatic nerve lesion model, this

implant supported axon regeneration and remyelination much better than silicone conduits in

comparison (Cai et al., 2005). Tubes consisting of collagen-coated PGA and braided PLA/PGA

were compared in dogs for 40 mm repair of the peroneal nerve. Regeneration, including recovery

of muscular function after 12 months, was best in PLA/PGA-collagen tubes, where it reached

80 % of the positive control (Ichihara et al., 2009).

Poly(lactic-co-glycolic acid)

With electrospinning, PLGA fibers were collected on a rotating mandrel, and tubular nerve

conduits from this material were tested as 10 mm bridges in the rat PNS. One month after

implantation, no inflammatory response was recorded but successful regeneration was observed

only in 5 of 11 rats (Bini et al., 2004). Substantial improvements were possible with combinations

of PLGA and chitosan, coated with CNTF. Histological evaluations showed that such an implant

Chapter I

32

was able to induce regeneration across a 25 mm tibial nerve gap. In a canine model, the conduit

was as good as the autologous nerve transplant (Shen et al., 2010).

Poly(ε-caprolactone)

Electrospun PCL fibers alone are good substrates for axonal regeneration and the migration of


33

various glial cells (Gerardo-nava et al., 2009). These properties are significantly improved via

blending with collagen (Schnell et al., 2007) or covalent binding with ECM-derived peptides

(Bockelmann et al., 2011, see below). Oliveira and coworkers seeded PCL conduits with

mesenchymal stem cells and tested them for repair of median nerves in mice. Motor function after

8 weeks was assessed with a grasping test in addition to histological examination. The study

demonstrated good regeneration and a positive effect of the implanted cells (Oliveira et al., 2010).

Poly(hydroxybutyric acid)

In experiments with a PHB conduit, excellent results were achieved when the tube was filled with

a GGF-containing gel (Mohanna et al., 2003, 2005). In a spinal cord implantation experiment,

PHB constructs showed good integration and axonal regeneration within the graft (Novikova et

al., 2008). These descriptions make clear that the best results were always obtained when polymer

fibers were functionalized or combined with ECM proteins (Chap.I 4.3).

4.2.3 Polyphosphoesters

In addition to various aspects of biocompatibility, which they share with poly(a-hydroxy

carboxylic acids), fibers of polyphosphoesters (PPE) offer the potential for coupling of bioactive

molecules. PPE were therefore developed as matrices for PN implants (Wang et al., 2001). Films

of electrospun fibers of poly(caprolactone-co-ethyl ethylene phosphate) (PCLEEP) served also as

a delivery device for NGF (Chew et al., 2005) or GDNF (Chew et al., 2007). While it was

possible to spin parallel aligned fibers of the PCL/PPE copolymer on a rotating drum,

incorporation of the growth factor in the spinning solution made this difficult. When rolled films

Figure 5 Evaluation of artificial nerve implants in animal experiments.

(a–c) Electron microscopy of transverse sections: (a) non-lesioned sciatic nerve of the rat; (b)

regeneration in an autologous nerve transplant; (c) regeneration in an implant with gradients of NGF

and laminin after 4 months; distance 10 mm from the transection site; Ax axons, M myelin sheaths, Sc

Schwann cell, scale bar 2 mm (Dodla and Bellamkonda, 2008). (d, e) Immuno-histochemical staining

of longitudinal sections of regenerated axons in an implant with electrospun microfibers in rat sciatic

nerve after 4 months: above neurofilament, axons; below S100, Schwann cells (Kim et al., 2008). (f)

Retrograde tracing with FluoroGold from the distal end of a chitosan/PGA implant showing axonal

regeneration of sensory neurons from the dorsal root ganglion in dogs at 6 months after implantation

(Wang et al., 2005). (g) Electrophysiological recordings to assess functional regeneration cross a gap

of 80 mm in dogs after 12 months. Collagen tubes with laminin-coated fibers were implanted into the

peroneal nerve. Traces are shown for compound muscle action potential (CMAP) of the anterior

tibialis muscle after stimulation of the sciatic nerve; motor evoked potentials (MEP) in the tibialis

muscle after electrical stimulation of the motor cortex; somatosensory evoked potentials (SEP)

recorded from the cerebral cortex after muscle stimulation; and spinal cord evoked potentials (SCEP)

with stimulation of the nerve stump distal to the graft (Matsumoto et al., 2000). (h) Functional

demonstration of regeneration (through the implants shown in d and e) in rats. In the grid walking

test the number of mistakes are counted while the rat walks across a grid. Implants with fibers

orientated in parallel (asterisk) achieved better results than implants without or with non-orientated

fibers, yet autologous nerve transplants were still better (Kim et al., 2008). (i) The largest nerve gaps

repaired with artificial implants, which showed recovery of motor functions, were found with dogs.

Note movement of the left hind limb (Toba et al., 2001)

Chapter I

34

of fibers were tested as a 15 mm nerve bridge in rats, nerve regeneration was observed in all

animals after 3 months, even without the growth factor. Electrophysiological examination showed

a beneficial effect of the encapsulated growth factor (Chew et al., 2007).

4.2.4 Poly(ethylene glycol)

This polymer, which is non-adhesive to cells, is used to produce gels and matrices or serves as a

linker for proteins. From poly(ethylene glycol), the star-shaped NCOpoly(ethylene glycol)-stat-

poly(propylene glycol) (sPEG) with a backbone of 80 % ethylene oxide and 20 % propylene

oxide has been synthesized. A range of biochemical functionalities such as isocyanate, acrylate,

and vinyl sulfone can be introduced to provide the basis for covalent modification. Using sPEG

with reactive isocyanate end groups, blends of PCL and PCLdiol were functionalized with ECM

derived peptides, which improved their properties as guidance substrates for axons and Schwann

cells from DRG (Klinkhammer et al., 2010; Bockelmann et al., 2011). Recently, an easy method

for peptide functionalization of polyester fibers via sPEG has been published, which is promising

for in vivo applications (Grafahrend et al., 2011).

4.2.5 Poly(2-hydroxyethyl methacrylate) and related hydrogels

Synthetic hydrogels, a prime example being poly(2-hydroxyethyl methacrylate) (pHEMA), have a

very high water content after polymerization. Gels consisting of poly(HEMA-co-methyl

methacrylate) (pHEMA-MMA) implanted in the spinal cord of adult rats integrated well. Axonal

regeneration into the implant was found by 8 weeks (Tsai et al., 2004). Hydrogels can be used to

make scaffolds with structured channels or to encapsulate functional molecules when growth

factors are added to pHEMA sheets during polymerization in aqueous solution. A nerve tube

consisting of porous pHEMA-MMA was tested in 10 mm rat sciatic nerve gaps. After 8 weeks,

the outcome was similar to autografts, and after 16 weeks 60 % of the artificial tubes were still

comparable to this positive control (Belkas et al., 2005b). In an experimental medical application,

pHEMA-MMA tubes filled with collagen resulting in a gel-like outer and porous inner structure.

As implants in the adult rat sciatic nerve, such conduits were very successful over a period of 8

weeks. In long-term survival studies, however, some of these implants tended to collapse and also

evoked inflammatory reactions as suggested by the presence of macrophages and giant cells in the

implants after 16 weeks (Belkas et al., 2005a). Copolymerization of HEMA with 2-aminoethyl

methacrylate provided primary amine groups for covalent binding of laminin-derived

oligopeptides. This approach was combined with physical modification of the gel, which

contained multiple longitudinal channels (Yu and Shoichet, 2005).

4.2.6 Oxidized polypyrrole

The electrically conducting polymer poly(pyrrole) (PPy) is interesting for neural applications

because electrical stimulation can influence axonal growth. In one study, electrospun PLGA fibers

were coated with PPy, and electrical stimulation of the conducting nanofibers enhanced neurite


35

outgrowth (Lee et al., 2009). This work and related pioneering studies (Gomez and Schmidt,

2007) were done with a PC12 cell line, however, this has only limited predictive value for

neurons in vivo. While the material properties of PPy, which was discovered in the 1960s, have

been thoroughly characterized, there is still limited experience with PPy as implant material in

vivo (Schmidt et al., 1997; Wang et al., 2004; George et al., 2009). In a tube of combined PPy and

silicone implanted to bridge a 10 mm gap in the rat, regeneration of nerve tissue was only

marginally better than that in the plain silicone tube (Schmidt et al., 1997; Wang et al., 2004;

George et al., 2009). Future studies will show whether the electrical conductance of PPy and other

conducting polymers (polyaniline, polythiophene) is a real advantage in PN implants in vivo.

4.3 Functionalization with ECM proteins and peptides

As noted above (Chap.I 3.3), tubular implants and scaffolds with longitudinal pores have been

made of collagen with very promising results. Similar success was obtained with artificial

polymers that were blended or coated with collagen and laminin (Inada et al., 2004, 2007;

Ichihara et al., 2009). On the other hand, synthetic polymers proved very useful in the

construction of nerve guides, not least because of their superior physical properties. For instance,

many hydrophobic polymers can be spun much more easily than ECM proteins (Klinkhammer et

al., 2009). Unfortunately, they lack the specific biochemical functions involved in nerve

regeneration (Chap.I 2.2 and 2.3). The purpose of functionalization is therefore to endow

synthetic polymers with molecular moieties that activate endogenous cells (neurons, Schwann

cells, endothelial cells) in a way that promotes axonal growth. Corresponding to the distinction

between secreted and surface-bound signals, two fundamental approaches have been employed:

binding of ECM molecules and controlled release of soluble growth factors.

4.3.1 Blends of synthetic polymers with proteins

The use of ECM molecules from the PNS is suggested because axons regenerate naturally along

the basal lamina. Thus, ECM proteins or peptides are coupled to synthetic polymers either by

blending, adsorption, or covalent binding (Koh et al., 2008) (Figure 6). This should promote cells

to migrate into the artificial implant and close the gap between the nerve stumps. As discussed

above, the main components of the basement membrane are the trimeric laminins plus collagen

IV. Fibrous collagens (I, III, V) are major components of the ECM (Koopmans et al., 2009).

Thus, a logical strategy was to blend synthetic polymers (e.g., of PLA with laminin (Koh et al.,

2008) or PCL with collagen (Zhang et al., 2005; Schnell et al., 2007)) and spin this material to

fibers, which could then be used for the fabrication of tubular implants or guidance structures.

PLA/laminin blends and PCL/collagen blends significantly improved the surface characteristics of

electrospun nanofibers in comparison with the pure synthetic polymers. Some groups were able to

spin fibers of pure ECM proteins (Chew et al., 2005; Heydarkhan-Hagvall et al., 2008); however,

Chapter I

36

some found it difficult to produce aligned fiber substrates of good quality with this method

(Klinkhammer et al., 2009).

4.3.2 Covalent functionalization of polymer fibers

Instead of whole ECM proteins, which are extracted from biological sources, specific peptide

sequences from laminin, fibronectin, or collagen can be synthesized (Wang et al., 2008; Dalton

and Mey, 2009). Their design was made possible by the knowledge of specific amino acid

Figure 6 Functionalization of polymer fibers for nerve regeneration.

(a) Functionalization with adsorption or covalent binding of ECM molecules to electrospun polymer

fibers (Koh et al., 2008). (b) Functionalization by blending of ECM molecules with synthetic polymers

before electrospinning (Schnell et al., 2007). (c) Functionalization by chemical bulk modification of

the electrospinning solution (Klinkhammer et al., 2010)


37

sequences in the ECM that bind to integrin receptors of Schwann cells and neurons. These

peptides are adhesive because of the intracellular connection of integrins to the cytoskeleton. In

addition, they activate a number of signal transduction cascades within the cells (Geiger et al.,

2001). So far, peptides with the amino acid sequences RGD, YIGSR, IKVAV, and longer

versions of those have been successfully applied. People from my laboratory showed, for

instance, that covalent binding of GRGDS to electrospun nanofibers of PCL improved their

ability to guide axonal growth (Bockelmann et al., 2011). Several functionalization strategies with

ECM molecules are also being tested in vivo: The interior surface of a chitosan tube was modified

with various laminin peptides. As an artificial bridge of the rat sciatic nerve, the tube with one of

these sequences promoted regeneration to a similar degree as a transplanted nerve, although not as

well as autologous implants (Wang et al., 2008). In addition to ECM proteins, the cell adhesion

molecules (CAMs, important examples for nerve regeneration are L1, N-CAM) constitute another

group of surface-bound signals that can induce axonal growth. These transmembrane proteins,

which are exposed on cell membranes and are not part of the ECM, are also being considered for

functionalization (Webb et al., 2001).

In conclusion, the chemical functionalization of synthetic materials with ECM molecules was an

important step in the development of artificial implants with biological activities. Apart from

these signals, which are always fixed to surfaces, peripheral nerve regeneration in vivo depends

on a number of signals that are secreted from Schwann cells or the neurons themselves (Chap.I

2.2).

4.4 Controlled release of growth factors

To apply soluble factors in vivo, the first ideas were to fill implanted devices with a solution of

the growth factor or with growth-factor-secreting cells. In a more sophisticated manner, carriers

are incorporated that gradually release the active molecules over longer periods. Several problems

have yet to be solved (de Ruiter et al., 2009). For instance, the biological activity should be

present for several weeks to months, especially in the case of longer nerve lesions. Also,

sterilization should be possible without compromising the function of the signals. As these issues

have been studied in several experiments, examples for the major paths of delivery are given

below. The reader can find additional references in (Siemionow et al., 2010).

4.4.1 Immobilization in the tubular wall

A common feature of artificial nerve conduits is an outer wall in the form of a tube. Different

techniques have been developed to immobilize growth factors such that they are released over

time from the scaffold wall to the inside. In one study, tubes were fabricated by dip-molding from

a solution of ethylene-vinyl acetate, bovine serum albumin, and FGF-2. In vitro assays revealed a

burst phase of FGF release for the first 3 days (50 % of total protein), which then declined to a

rate of 0.1 – 0.5 % per day. In vivo, after 4 weeks these tubes were clearly superior in bridging a

Chapter I

38

15 mm sciatic nerve gap compared to tubes without the growth factor (Aebischer et al., 1989).

Madduri and coworkers incorporated NGF and GDNF in conduits consisting of collagen

(Madduri et al., 2010a) or fibroin (silk) fibers (Madduri et al., 2010b). During biological

degradation of the scaffold, the neurotrophic factors were gradually released.

4.4.2 Growth factors released from gels

Advanced conduit designs involve internal scaffolds for stabilization and guidance of the

regenerating tissue. Hence, an obvious approach is to incorporate the growth factor into this

polymer by mixing before gelling. This was done with NGF and NT-3 in pHEMA (Moore et al.,

2006). The growth factor can be equally distributed over the conduit or in the form a gradient to

attract continued regeneration in the distal direction of the implant. In some cases, it could

become a problem that axons are trapped within the chemoattractive environment of the conduit

(“candy store phenomenon”). In a successful example of the gradient strategy, 40 mm gaps in the

peroneal nerve of rabbits were bridged within little more than 6 weeks. This was achieved using

PHB conduits filled with GGF (pro-neuregulin-1), which was suspended in alginate (Mohanna et

al., 2005). Neither empty nor alginate-filled tubes without GGF were able to support growth of

Schwann cells or axons across the complete gap, even after 9 weeks (Mohanna et al., 2003).

Integration of polysaccharides into the polymer can enhance growth factor retention due to

hydrophobic, electrostatic, or ionic interaction. This is a specific activation effect with heparan

sulfate proteoglycans and growth factors FGF-2, VEGF, and GDNF (Barnett et al., 2002; Forsten-

Williams et al., 2008). This synergistic interaction was exploited by using alginate/heparin gels as

a matrix for FGF-2, which increased vascularization and faster axonal growth through the

artificial implants (Ohta et al., 2004). Combining growth factors with ECM proteins was also

promising. Dodla and colleagues devised an agarose gel to apply gradients of NGF and laminin.

Both gradients had a positive effect on nerve regeneration (Dodla and Bellamkonda, 2008).

4.4.3 Microspheres as delivery vectors

Instead of inducing growth factors into the conduit wall or the inner matrix, they can be

encapsulated in microspheres before incorporation. With this approach, the release kinetics can be

tested and tuned beforehand in vitro. Microspheres of PLGA and PLA with GDNF were

embedded into the inner half of a PCL-tube fabricated by dip-coating. The microspheres were

formed by adding an oil-in-oil emulsion of PLA and PLGA (the latter containing GDNF) drop-

wise to an aqueous solution, followed by centrifugation and lyophilization. The GDNF release

kinetics observed in vitro showed a strong release of 4.6 ng GDNF per mg microsphere during the

first day. The subsequent release approached zero-order kinetics and reached a cumulative release

of 6.4 ng/mg microsphere at day 64. The bridging of a 15 mm rat sciatic nerve gap with this nerve

conduit was much better than without the GDNF, similar to results obtained with nerve autografts

(Kokai et al., 2011). In another study, nerve conduits containing PPE microspheres with NGF

5 Conclusion

39

were successfully used to bridge a 10 mm rat sciatic nerve gap. Three months after implantation,

rats with NGF-microsphere conduits showed a higher percentage of positive reflex response than

controls after pinching distal nerve trunks (Xu et al., 2003).

4.4.4 Release from electrospun fibers

Aligned electrospun fibers, which are perhaps the most promising topological cues for axon

guidance, were also used to deliver growth factors. For example, PCLEEP fibers containing

GDNF have been obtained by mixing GDNF into the polymer solution followed by

electrospinning. The fibers were spun into a PCLEEP film, which was later rolled to form a tube

on the inner surface of a nerve guide (Figure 2g). Release kinetics were characterized by an initial

burst of about 30 % of protein followed by relatively stable release until leveling off after almost

2 months. In vivo investigations with 15 mm rat PN lesions revealed a significantly higher

number of myelinated axons inside the GDNF-implant compared to the controls (Chew et al.,

2007). Recently, a conduit was designed that consisted of a PLA-co-PCL shell made of

electrospun nanofibers filled with an NGF-containing core of bovine serum albumin. Again, when

tested as a sciatic nerve bridge of 10 mm in rats, results after 12 weeks were similar to those

obtained with autografts (Liu et al., 2011).

In summary many different and equally promising strategies of functionalization have been

followed during the last decade (Figure 5 and Figure 6). Unfortunately, since there are hardly any

studies that compare different strategies of functionalization, at this moment it is difficult to

decide which molecules should be preferred in the future and which delivery strategy is most

suitable.

5 Conclusion

In contrast to long-distance connections within the CNS such as the optic nerve or the white

matter tracts of the spinal cord, lesioned peripheral nerves contain the inherent cellular and

molecular components that are required to support successful axonal regeneration. The

development of axon growth-promoting devices that incorporate a number of these components

has become an expanding area of research. Although a number of biocompatible materials have

been developed, which offer great potential for therapies of peripheral nerve injury, several

technical or scientific challenges need to be addressed.

5.1 Standardized control of efficiency

Many experiments carried out using different biomaterials are not comparable with each other

because different cell cultures, animal models, or criteria of success were used. In general, I have

observed that a considerable number of publications reporting innovative materials failed to use

appropriate biological tests. Papers falling into this category have not been discussed in the

present review. Even in many animal studies, the inflammatory reactions, effects on gene

Chapter I

40

expression, and long-term results are rarely investigated. Clearly, a systematic approach with

internationally agreed standards would be desirable for such investigations to allow a better

degree of comparison.

5.2 Structured three-dimensional implants

So far, most studies of PN regeneration in vivo have only implanted hollow tubes, where axons

start to grow on the inner surface of the implant. Thus a major technical challenge is the

construction of three-dimensional nerve bridges with an internal architecture that is capable of

guiding axonal growth and migration of the endogenous Schwann cells. In my opinion, the most

promising approaches are the three-dimensional integration of microfibers in parallel orientation

and the production of orientated channels in gels or matrices of ECM proteins.

5.3 Biomimetic functionalization of implants

With the use of proteins and peptides derived from ECM, artificial implants can already mimic

the signals that regenerating axons encounter in the environment of the lesioned PNS. However,

the concept of targeted use of factors that specifically activate Schwann cells, sensory neurons, or

motor neurons is still in its infancy. In addition to the selection of the appropriate combination of

molecular signals, the ideal temporal and spatial presentation of such molecules for the efficient

support of axon regeneration and tissue repair remains unknown. In this context, it is possible that

neurotrophic signals might be required for several months and with increasing gradients towards

the distal end of the implant.

5.4 The thirty millimeter mark

Studies have demonstrated that peripheral nerve regeneration through artificial implants can cover

distances of up to 30 mm but rarely beyond. Successful motor fiber regeneration appears to be

more difficult than that of sensory axons, and thin fascicles regenerate more easily than large

caliber nerves. Since larger gaps must occasionally be bridged in cases of human nerve injury, the

usefulness of future constructs will be judged by their ability to support motor and sensory axon

regeneration over long distances. For this, it is likely that experiments with animals larger than the

rat will be required. I am convinced that results from basic research on mechanisms of axonal

growth in combination with engineering methods (such as electrospinning, freeze-drying, and

chemical functionalization) will result in further progress, with the development of artificial nerve

implants that approach or even exceed the efficiency of the autograft. Some very substantial

advances have already been made in the last few years. Thus, there is great hope and expectation

that such strategies will provide the neurosurgeon with a range of cell-free scaffolds that can be

used on demand.

1 Introduction

41

Chapter II

Manufacture and characterization of three-dimensional fiber arrays

and assessment of guidance properties in vitro

1 Introduction

Peripheral nerve (PN) lesions can be surgically repaired by suturing the proximal and distal nerve

stumps together. In these cases, axons regenerating from injured motor, sensory, and autonomic

neurons are able to grow within the scaffolds of Schwann cells and extracellular matrix (ECM) of

the injured nerve (Dalton and Mey, 2009). When PN injuries result in larger gaps between

transected nerve ends, these have to be bridged, for instance, by transplanting a sensory nerve

from the same patient. Unfortunately, this approach is functionally successful only in about half

of the patients and has other disadvantages, such as sensory loss at the donor site and a limited

availability of transplant material (Deumens et al., 2010). Currently, much research is therefore

devoted to developing artificial nerve conduits that can replace autologous transplants (Deumens

et al., 2010; Siemionow et al., 2010; Mey et al., 2012).

For clinical use in humans, only simple tubular conduits have so far been approved in the United

States and Europe (Siemionow et al., 2010). The majority of animal experiments are also based on

the design of the hollow tube. Because axons grow on surfaces and must be guided toward their

peripheral target, better results are expected if internal structures could support axonal growth and

induce colonization with host-derived Schwann cells, which are required for trophic support and

the subsequent remyelination of regenerated nerve fibers (Hoffman-Kim et al., 2010). Recent

advances in the design of artificial nerve implants therefore focus on topographical features that

guide axonal regeneration and cell migration. Basic design strategies for nerve bridges with

internal guidance structures are based on (1) large interior channels to separate individual fiber

fascicles (Clements et al., 2009; Yao et al., 2010), (2) scaffolds with many longitudinally

orientated pores (Ceballos et al., 1999; Möllers et al., 2009), or (3) parallel aligned polymer fibers

produced by melt extrusion (Matsumoto et al., 2000; Ribeiro-Resende et al., 2009) or

electrospinning (Teo and Ramakrishna, 2006; Kim et al., 2008) (Figure 7).

The approach presented here is based on electrospinning of the biodegradable polyester poly(ε-

caprolactone) (PCL) using two separated target electrodes, such that parallel fibers accumulate

between them (Schnell et al., 2007; Klinkhammer et al., 2009). The resulting microfibers have

been shown to be excellent growth substrates for glial cells and axons of sensory neurons

(Gerardo-nava et al., 2009; Bockelmann et al., 2011). Given successful in vitro tests of several

two-dimensional growth substrates a major challenge in scaffold design consists in the

construction of three-dimensional (3D) structures that can be implanted in vivo. In this study, I

describe a technique for 3D spinning of parallel microfibers, their embedding in collagen gels,

Chapter II

42

and incorporation in biodegradable PCL tubes. The biocompatibility of the resulting 3D scaffold

was investigated using dorsal root ganglia (DRG) explants.

2 Material and Methods

2.1 Electrospinning of three-dimensional fiber arrays

Solvents and PCL used for electrospinning were purchased from Sigma-Aldrich Chemie GmbH

(Munich, Germany). Electrospinning was performed with 8 wt% and 12 wt% PCL (Mn = 70,000 –

90,000 g/mol) dissolved in chloroform/methanol (3:1). Unless otherwise indicated, data are

reported for fibers made from 12 % PCL. The spinneret consisted of a syringe with a 20G flat

tipped needle connected to a metal disc with 90 mm diameter. After various target electrodes had

been tested, the standard target configuration consisted of two opposing aluminum bars with a

distance of 18 mm between them. These were mounted on a tapered base, which was connected

with its narrow end to a grounded metal disc (Figure 8a). Of the different target devices, all made

Figure 7 Conduit design for artificial nerve guides. (a) Empty tube used for bridging gaps between proximal and distal nerve stumps,used in most experiments so far; (b) tubular implant with multiple intraluminal channels (e.g. de Ruiter et al., 2008; Yao et al., 2010); (c) tube filled with an internal matrix structured by longitudinally oriented pores or channels (e.g. Bozkurt et al., 2012); (d) aligned polymer fibers as longitudinal guidance structures (e.g. Kim et al., 2008; Ribeiro-Resende et al., 2009); (e) present construct consisting of 3D aligned PCL fibers embedded in layers of collagen gel.


43

of aluminum (Figure 8b–e), a stair case-like target was used for spinning of 3D fiber arrays

(different sizes were successfully tested, e.g., 12 steps, 20 mm breadth, 1 mm step size; five steps,

13 mm breadth, 1 mm step size; (Figure 8d,e and Figure 9a). Standard spinning conditions were

17 kV, with 20 cm distance to the collector without mechanical ejection of the polymer solution

by a syringe pump resulting in a feed rate of about 0.5 ml/h. However, depending on ambient

conditions the distance between spinneret and target was adjusted between 17 and 20 cm to give

Figure 8 Electrospinning setup for the collection of 3D fiber arrays. (a) Schematic diagram of the spinning device consisting of a syringe, 20 gauge needle connected with a metal plate (spinneret), high voltage power supply, and grounded collection target, which is mounted on a conducting v-shaped base likewise connected to a metal plate. (b–e) Photos of different target designs to illustrate accumulation of parallel fibers. Two parallel bars (b) resulted in a single layer of parallel fibers accumulating on the upper edge. Substituting edged bars with a curvature allowed (c) accumulation on a slightly wider depth on the target. The best target for 3D-spinning is characterized by a stepped design (d) resulting in a sequential accumulation of fibers from bottom to the top edge (e, arrows).

Chapter II

44

Figure 9 Three-dimensional arrays of parallel fibers. (a) Electrospun fibers suspended in a step target are shown from the side. (b) Projection of a 400-mm confocal scan from above demonstrates the parallel alignment of the fibers. (c) SEM image of fibers with 200-fold and 5000-fold magnification. High magnification photos were used to determine the fiber diameters (fiber seen as a vertical line in the inset; 808 ± 175 nm, mean ± SD)

best results. For spinning of one batch of fibers I applied the voltage for approximately 30 s to the

spinneret. To handle the three-dimensional array of fibers after spinning, the bars were attached to

a U-shaped frame before disassembling it from the base. To visualize electrospun fibers with

fluorescence microscopy, the lipophilic dye DiI [Molecular Probes, D3911, 75 ml DiI/DMSO

(1/1000) in 20 ml] was added to the PCL solution.

2.2 Physical characterization of fiber arrays

Electrospun fibers suspended on the targets were examined with a Leica TCS SP2 confocal

microscope (Leica Microsystems GmbH, Wetzlar, Germany). For documentation of alignment

and density of the fiber arrays image stacks were produced from the upper most fiber plane

reaching 2 mm down into the lowest fiber plane. The image stacks were processed for further

analysis using Image J software. For evaluation of fiber density, I divided a stack into five bins of

400 µm, made a projection of all images in one bin and counted the number of fibers crossing the

midline. For assessing fiber alignment, the mean direction of all fibers in a stack was determined

and the deviation angles of the individual fibers were plotted. The standard deviation (SD) of

deviation angles was calculated using MATLAB. Fiber diameter was determined by sputter

coating the samples with 30 nm gold and

performing scanning electron microscopy

using a FEI/Philips XL30 FEG ESEM

with a 10 kV electron beam. Using high

magnification (5000×) images of single

fibers I averaged the thickness of each

fiber from five positions within an image.

Mean fiber thickness and standard error

(SEM) were calculated from 23 fibers of

five different samples.


45

2.3 Inserting the 3D fiber arrays into a conduit

For biological testing, the electrospun fibers were inserted in gels that consisted of type I

collagen. For the matrix, I prepared a 2.4 mg/ml solution of rat tail collagen I (Life Technologies,

A1048301) according to the manufacturer’s protocol. In brief, one part 10× basal medium Eagle

was added to eight parts rat tail collagen I solution. I neutralized the pH with 1 M NaOH and then

filled up to 10 parts with distilled water (DRG cultures) or with cell suspension to obtain

106 cells/ml. To prevent premature gelation, the collagen solution was kept on ice during the

whole procedure. To produce composites of collagen and PCL fibers, a drop of 50 ml collagen

solution was applied onto a poly-L-lysine coated cover slip. Using a U-shaped frame to hold the

target electrodes the fiber array was deposited into the collagen solution and detached from the

frame using hot forceps. To induce gelation, the composites were placed in an incubator for 30 –

40 min at 37°C and 95 % humidity. This step could be repeated several times to incorporate

additional layers of fibers.

For manufacturing implantable tubes to be filled with the composite of gel and fibers, porous,

biodegradable PCL tubes were made by dip coating. Crystalline NaCl was ground to fine powder

in a mortar and was mixed with an equal amount (wt) of a 12 % PCL-solution in

chloroform/methanol (3:1). Steel wires of 2 mm diameter were repeatedly dipped into the

NaCl/PCL suspension to produce conduits of 1 – 2 cm length and 500 µm wall thickness (for a

final construct of 1.5 cm). The surface was leveled after 1 min drying at RT. After another 5 h of

drying, the conduits were removed from the wires and transferred to a Petri dish with stirred

deionized water for 24 h to dissolve the NaCl(Chang, 2009). Water was changed twice. Conduits

were cut lengthwise in half and filled with gel/fiber stacks as described above for glass cover

slips. Subsequently, half-tubes were resealed using the PCL solution and equilibrated with cell

culture medium. All steps were performed under sterile conditions. The thickness of the tube wall

was determined with scanning electron microscopy.

2.4 Cell and organ culture

All cell culture reagents were purchased from Life Technologies GmbH (Darmstadt, Germany).

Before the experiments, astrocytoma cells (U373) were grown in Dulbecco’s modified Eagle

medium (DMEM) with Glutamax and enriched with 10 % fetal calf serum (FCS) and 0.2 %

penicillin/streptomycin at 37°C, 5 % CO2. The culture medium was changed every second day

and cultures were split twice a week. For the experiments, cells were washed with phosphate

buffered saline (PBS) and detached from the surface of the culture flask using 0.25 %

trypsin/ethylenediaminetetraacetic acid (EDTA). After trypsination was stopped by adding 5 ml

serum containing medium, cells were centrifuged and resuspended in culture medium. Cell

numbers were determined using a Thoma chamber.

Chapter II

46

For organ cultures, I dissected DRG from 10-day-old chicken embryos. The embryos were

decapitated and the skin removed. Accessing from the dorsal side, sharp forceps were inserted

close to the lumbar vertebral column. By opening the forceps the hind limb could then be

separated from the torso. The DRG, remaining on the inside of the thigh, were removed and

collected in cold, sterile PBS, where they were kept until being transferred to the collagen matrix

(< 30 min) and before adding of the PCL-fibers. After gel formation, the samples were placed in a

24-well plate with 1 ml culture medium and incubated at 37°C and 5 % CO2 for six (single cells)

or seven days (DRG). The medium was changed once after three days.

The biological compatibility of the complete conduits was also tested with DRG cultures. After

coating the inner side of the longitudinally sectioned half-conduits with poly-L-lysine, several

layers of collagen and PCL-fibers were added together with a DRG explant. To allow microscopic

evaluation, I sealed the surface with a poly-L-lysine coated cover slip and transferred the

assembly upside down to a culture well, where culture media could only enter through the wall or

the endings of the conduit. After 30 min gelation at 37°C I added 1 ml of culture medium to the

well. The culture was maintained for seven days with medium change once after three days in

vitro. Axonal growth and cell migration was investigated in whole mounts.

2.5 Immunocytochemistry

Samples were fixed 1 h with cold 4 % paraformaldehyde/phosphate buffer, rinsed with PBS and

blocked with 1 % bovine serum albumin, 4 % normal goat serum and 0.1 % TritonX-100 in PBS

for 1 h. Incubation with the primary antibody [either 1/500 rabbit-anti-S100 (Sigma-Aldrich,

Munich, Germany, S2644) or 1/500 rabbit-anti-NF200 (Sigma-Aldrich, N4142)] was performed

over night at 4°C in blocking solution. Incubation with secondary antibody, Alexa Fluor 488 goat-

anti-rabbit (Life Technologies GmbH, Darmstadt, Germany, A11008), 1/500 in blocking solution,

was done for 1 h at room temperature followed by nuclear staining with 4’,6-diamidino-2-

phenylindole (DAPI, 1:5000) for 10 min. Negative controls were performed without primary

antibodies. Samples were placed upside down on object slides using Fluoprep (Biomerieux) to

stabilize fluorochromes.

2.6 Fluorescence microscopy

Images of the U373 astrocytoma cell cultures were captured using a Leica TCS SP2 confocal

microscope and were further processed using Image J software. Cell migration and axonal

outgrowth from DRG were analyzed with a Leica DM5000B epifluorescence microscope using

filter cubes L5, N2.1 and A4 for Alexa Fluor 488, DiI and DAPI nuclear staining. Photos were

recorded with a Leica DFC350 FX digital camera. Composite pictures of entire DRG and all

outgrowing cells and axons were assembled by means of a Leica DMI6000 microscope with

LASAF software. Individual cells and their interaction with PCL fibers were imaged using Leica

TCS-SP5 confocal microscopy and further processed using Image J.

3 Results

47

2.7 Quantification of Schwann cell migration and axonal growth

On composite photographs of DRG four tangents were drawn to the edges of the DRG in parallel

and orthogonal to the fiber direction or, in case of the controls, horizontal and vertical with

respect to the photographic frame. For cell migration, I measured the distance between the farthest

cell body and the corresponding tangent. For axonal outgrowth, I counted the number of visible

axonal endings and measured the distance between the 10 % longest axons and the corresponding

tangent. In the statistical analysis the individual DRG were considered as independent samples.

Conditions were compared using Student’s t-test, calculated with MATLAB.

3 Results

3.1 Electrospinning of parallel fibers in a 3D configuration

The configuration of the fiber array produced with electrospinning depends on the geometry of

the target which is used. Although a simple plate target results in a mat of non-oriented fibers,

parallel fiber alignment can be obtained with targets consisting of a rotating drum or two parallel

bars. Using the latter design, I collected PCL fibers that were suspended in the air between the

two target bars (Figure 8a). Fibers were aligned in parallel but formed one layer as they attached

to the upper edge of the bars (Figure 8b). To achieve a 3D array of the accumulating fibers,

various target configurations were tested by modifying the bars. For instance, replacing the sharp

edges with a rounder curvature led to the accumulation of fibers on a slightly wider area of the

target, yet the distribution was not easy to control and was frequently different between the left

and right bar (Figure 8c). Taking advantage of the fact that the fibers tend to accumulate on edges,

where electric field lines converge, I designed step-like targets with multiple edges facing each

other (Figure 8d). On this kind of target, the accumulation of fibers starts at the lowest edges.

With increasing density as fibers accumulate on these edges, electrostatic repulsion between them

forced further fiber accumulation towards the next higher steps. In this way, a consecutive

accumulation of fibers from bottom to top was achieved (Figure 8e) resulting in a 3D

configuration of PCL fibers. A view from the side of a 3D fiber array spun onto the final version

of the target is shown in Figure 9a. Microscopic evaluation of fluorescently labeled fibers (Figure

9b) and SEM visualization (Figure 9c) demonstrated the degree of their alignment.

Chapter II

48

Figure 10 Quantitative characterization of the 3D fiber array. (a) Fiber density was determined from sequential confocal scans of 400 mm thickness through the fiber arrays suspended in the target frame (mean ± SEM, n = 3 targets). The fiber density decreases from top to bottom with the bottom layer still containing more than 18 fibers/mm2. (b) The histogram depicts variability of fiber orientation as individual deviation from the mean orientation (n = 559 fibers from six samples. Angular standard deviation is 2.46°).

3.2 Characterization of electrospun 3D PCL-fiber arrays

To quantitatively characterize the fiber arrays, the fiber density at different depths inside the array

was binned (bin 1 to 5; Figure 10a). The fiber density decreased from the top to the bottom of the

3D array, with the highest and lowest fiber density, 86.7 ± 5.1 fibers/mm2 (area of cross section, n

= 5 fiber stacks, mean ± SEM) and 18.3 ± 6.4 fibers/mm2

in the first and last bin, respectively

(ANOVA p < 0.005; F = 8.00; df = 4). The mean fiber density was 48.8 ± 4.5 fibers/mm2.

To assess fiber alignment, I determined the direction of the individual fibers, calculated the mean

orientation of all fibers and for every fiber its deviation angle from the mean orientation (Figure

10b). The alignment of electrospun fibers was high, as fiber angles showed a narrow Gaussian

distribution with around 70 % of the fibers deviating less than 2.5° from the mean fiber

orientation (n = 559 fibers; SD = 2.46°). Only about 1 % of all measured fibers deviated more

than 10° from the mean. The maximum deviation was 15°.

Fiber diameters were determined on the basis of scanning electron microscopy photos (Figure 9c).

As described previously (Jha et al., 2011), fiber diameter was dependent on the PCL

concentration of the spinning solution used. A concentration of 8 % PCL resulted in a fiber

diameter of 660 ± 190 nm (mean ± SEM). By raising the concentration to 12 %, the fiber diameter

was increased to 808 ± 175 nm (p < 0.05; compared by student’s t-test).

3.3 Suspension and manipulation of 3D fiber

arrays in biocompatible gels

The 3D arrays of PCL fibers were embedded into a

collagen matrix by inserting the fibers into the gel

(Figure 11a). To handle the 3D fiber stacks for this

purpose, electrodes were fixed to a metal frame

(Figure 11b) and detached from the spinning setup

with the suspended fibers between them. The

embedding caused some compression of the array

thereby increasing the fiber density.

3 Results

49

To achieve larger 3D stacks that were not limited by the size of the electrospinning targets, this

procedure could be repeated several times by stacking alternate layers of fiber arrays (Figure 11c–

e). The thickness of the gel layers between 3D fiber arrays was determined by the viscosity of the

gel solution. To reduce displacement of the fibers by the collagen solution it was important to

introduce them before gelation of the collagen took place. Gelation was initiated by heating the

collagen to 37°C, but to some extent changes in viscosity of the collagen already occurred at room

Figure 11 Embedding of 3D fiber stacks in collagen. (a) Schematic drawing of the embedding process. The collagen solution is applied onto a poly(lysine)-coated coverslip. Subsequently, fibers are dipped into the collagen solution and detached from the frame using hot forceps. Gelling was induced by incubation of the composite at 37 C for 30 min. To add fiber layers, the process was repeated several times. (b) For handling of the fiber stacks a U-shaped frame was used to keep the two parts of the target in place when detached from the base. (c–e) Arrangement of fiber stacks in the collagen gel shown in cross section after one (c), two (d), and 10 (e) rounds of deposition. (f) One single 3D array of PCL fibers was maintained in a gelatin gel which was cast over the fibers suspended between the target electrodes.

Chapter II

50

temperature. Thus with collagen gels I obtained a matrix with several 3D arrays of fibers

separated by regions devoid of fibers (Figure 11d and Figure 7e). A standard array now consists

of 10 layers of fibers and gels with a distance of 70 – 120 µm between the layers (Figure 11e)].

An alternative method to contain the 3D fiber array within a hydrogel consisted in casting the gel

solution over the fibers (Figure 11f)].

The effect of fiber-containing collagen matrices on cell shape was first tested using the glial cell

line U373. These cells grew well within the scaffolds and produced cellular processes. When the

cells were cultivated inside the 3D collagen matrix without PCL fibers, I observed a more diverse

morphology than in 2D cultures. In the 3D gels, many cells had complex cell bodies with several

processes in various directions (Figure 12a). When PCL fibers were added in 2D or 3D cultures

cell bodies tended to attach to the polymer fibers (Figure 12b). In 3D collagen matrices, all cells

Figure 12 Gel-embedded PCL fibers affect cell morphology. To visualize the effects on cell shapes U373 cells were cultivated for 6 days in a 3D collagen matrix with and without PCL fibers. (a,b) Confocal microscopy images show immunostaining against S100 (green) and nuclear staining (DAPI, blue). PCL fibers were stained red using DiI. Panels (a) and (c) show cell morphologies in the absence of PCL fibers, panels (b) and (d) with PCL fibers in the gel. Different magnification are shown in (a) and (b) with all scale bars measuring 50 mm. (c,d) Reconstructions of cells from both conditions to display the different cellular morphology in space. The red arrow in (d) points at the PCL fiber.

3 Results

51

which made contact with PCL fibers produced only two processes pointing in opposing directions

along the path of the PCL fiber. Additional processes could only be observed when cells had

contact with more than one fiber. Despite their ability to grow in complex shapes within the 3D

collagen gels (Figure 12c) the presence of PCL fibers imposed a fusiform morphology on the cells

(Figure 12d).

3.4 Influence of PCL fibers on axonal growth

To assess the influence of the 3D gel/fiber substrates on Schwann cells and sensory neurons, DRG

were explanted on the gels, and axonal growth and glial cell migration were quantified.

Representative images of DRG cultures in collagen gels with and without PCL-fibers are shown

in Figure 13. The 3D gel/fiber matrices caused a directional outgrowth of axons (Figure 13a) and

a preference in the direction of cell migration (Figure 13b).

When comparing the total number of axonal endings I found no significant difference between the

two conditions. Although initial neurite outgrowth was not influenced by the PCL fibers, this was

the case for axonal elongation. Under control conditions, when no fibers were contained in the

gel, the statistical analysis revealed no preference of axonal growth in any direction (Figure 13a)

and Figure 14a). At 4 DIV and 7 DIV the ratio of axon length in two orthogonal directions of the

Figure 13 Gel-embedded PCL fibers as a guidance structure for Schwann cells and sensory neurons. Dorsal root ganglia were cultivated for 4 and 7 days in a 3D collagen matrix with and without PCL fibers as control. Axons were immunostained against NF200 (a) and nuclei were stained with DAPI (b). Yellow arrows indicate the orientation of PCL fibers in the samples.

Chapter II

52

image was close to one (Figure 14b). In contrast, collagen gels that contained aligned PCL-fibers

had the effect that axons in direction of the PCL fibers were significantly longer than in the

absence of fibers, and axons in the direction orthogonal to the PCL fibers were significantly

shorter (Figure 14a), t-test p < 0.01 at 4 DIV]. For instance, after 4 DIV mean axon lengths in

parallel with PCL fibers was 1148 ± 81 µm, whereas in the orthogonal direction they only reached

378 ± 36 µm.

As axonal growth along the electrospun fibers continued in the cultures, the difference between

the two directions was even larger at 7 DIV. Axons also continued to grow in collagen gels

without PCL fibers. For this reason the difference in axon length between gels without PCL fibers

and the preferred direction in the collagen/PCL fiber matrix was no longer significant. However,

the orientation index demonstrated quantitatively that the fiber containing matrices continued to

Figure 14 Quantification of the guidance properties of gel-embedded PCL fibres. (a) Axon length in different directions is plotted as a result of the different 3D substrates. Data show mean of the 10 % longest axons of every explant. Substrates were collagen gels without or with PCL fibers. For the controls axonal growth was analyzed in arbitrarily defined horizontal and vertical (h and v) directions; for fiber containing gels axon lengths was assessed in parallel and orthogonal to the fiber direction (p and o). (b) An orientation index for axon length was determined as the ratio between the two directions measured for every condition (h/v or p/o). (c) The distance of cell migration from the DRG was quantified in a similar manner, and (d) an orientation index was calculated accordingly. Asterisks indicate significant differences between the PCL fiber sample and control; pound symbols indicate differences between the two directions within PCL fiber samples. Bars show mean ± SEM (n ≥ 6 DRG explants; Student’s t-test; p* < 0.05; p*** < 0.001; p### < 0.001).

3 Results

53

exert a strong guidance effect on the growing axons, whereas outgrowth was random in isotropic

collagen gels (Figure 14b). Comparing the results at both time points indicated that under all

conditions axonal elongation slowed down during the time course of cultivation.

3.5 Influence of PCL fibers on glial cell migration

The data for cell migration demonstrated a similar effect of the aligned PCL fibers (Figure 13b

and Figure 14c,d). Under the control condition of 3D gels without fibers there was no significant

difference in the distance of cells that migrated from the DRG in vertical or horizontal direction

(e.g., 4 DIV: 834 ± 52 µm vs. 780 ± 60 µm). When cultivated in the presence of aligned PCL-

fibers, cell migration parallel to the fiber orientation reached significantly longer distances (e.g.,

at 4 DIV: 1509 ± 55 µm) in comparison with cell migration perpendicular to the PCL fibers

(4 DIV, 7 DIV: both p < 0.001) and also in comparison with cell migration in 3D gels without

fibers (4 DIV: p < 0.05; 7 DIV: p < 0.001). As for axonal growth the orientation index was highly

significant at 4 DIV and 7 DIV (Figure 14d).

3.6 Incorporation of gel/fiber composites in a tubular conduit

To use the gel/fiber composite scaffold as a substrate for PN regeneration I inserted it into a PCL-

conduit, which was manufactured by dip-coating a steel rod with PCL solution. The addition of

NaCl to the PCL-solution and subsequent leaching resulted in the formation of a tubular conduit

wall with a porous surface (Figure 15). It also increased the flexibility of the conduit and hence

the suitability for in vivo applications. Surfaces and cross sections of tubes were examined with

scanning electron microscopy (Figure 15a,b), and parameters of the dipping technique

(concentration of the solutions, repetitions of coating) were adjusted according to the results.

PCL tubes were longitudinally cut in half, gel/fiber scaffolds were inserted into the half tubes, and

artificial nerve bridges were then produced by sealing two half tubes with PCL-solution (Figure

15c). To investigate the biocompatibility of the constructs, DRG explants were cultivated inside

the half conduits, which were sealed with a cover slip, so that exchange of metabolic matter could

only occur through the gel and PCL wall or via the endings of the conduit. The cultivated DRG

showed no visible change in appearance compared with those cultivated in open 3D gels. After

7 DIV, axon growth and cell migration had occurred within the 3D matrix (Figure 15d) and on the

surface of the cover slips. Anti-NF200 immunohistochemistry revealed widespread axonal growth

and similar axon lengths to those observed in 3D cultures that were not enclosed by tubular

conduits. Nuclear DAPI staining of cells that had migrated out of the DRG indicated cell

migration was also not impaired.

Chapter II

54

4 Discussion

To develop artificial scaffolds for PN regeneration I have described a method for the 3D assembly

of electrospun PCL fibers in parallel alignment. Consecutive embedding of these fibers in layers

of collagen gels and incorporation in a tubular conduit resulted in a 3D scaffold for axonal

guidance, whose biological use was tested with DRG explants. Schwann cells and sensory axons

adhered to the PCL fiber substrate and were guided by the orientated structure inside the gel/fiber

scaffold. Within the 3D collagen gels contact with the microfibers caused the cells to assume a

spindle-shaped morphology following the fiber direction.

Previous developments of fiber substrates revealed two major challenges with respect to the goal

of an artificial nerve implant, (1) the incorporation of microfibers in a 3D configuration, and (2)

the functionalization of the growth substrate to attract axonal growth over long distances

Figure 15 Integration of the composit scaffold into a biocompatible conduit. Conduits were produced from a mixture of a 12 % PCL solution and NaCl by dipmolding and subsequent leaching of the NaCl fraction. (a,b) SEM images depict the porous structure of the conduits wall (cross-section in a) and surface (b). (c) Scheme of the culture model used to verify biocompatibility of the conduit. In the inset a photograph of the collagen-fiber scaffold integrated into the tubular PCL conduit is shown. (d) Axonal growth and cell migration occurred within the structured 3D conduit. Axons are marked with NF200 immunostaining, migrated cells with nuclear DAPI stain.

4 Discussion

55

(Bockelmann et al., 2011). The present study contributes to solving the first of these problems and

provides a 3D platform to address the second.

4.1 Three-dimensional configuration of guidance structures

Artificial nerve bridges are designed as tubular grafts that connect the proximal with the distal PN

stump (Siemionow et al., 2010; Mey et al., 2012). When hollow cylinders with a single lumen are

used, regenerating axons and migrating cells accumulate on the inside of the tubular walls.

Constructs with several lumens were fabricated to increase that surface area and to separate

different fiber bundles within the conduit (de Ruiter et al., 2008; Clements et al., 2009; Yao et al.,

2010). However, increasing the number of compartments with these methods, for example by

using seven channels or by incorporating three films, did not improve regeneration (Clements et

al., 2009; Yao et al., 2010). Although implants have also been developed that were filled with

non-orientated “isotropic” matrices (Labrador et al., 1998; Toba et al., 2001; Inada et al., 2004;

McGrath et al., 2010), most studies indicate that orientated 3D structures with a large surface area

are required to provide an adequate substrate for long distance axonal guidance (Hoffman-Kim et

al., 2010; Mey et al., 2012). I can distinguish two basic designs that provide such internal

guidance structures within tubular implants (Figure 7). The tubes contain either longitudinal fibers

or a porous matrix with longitudinal orientation of the pores. Conduits that consist of polymer

fibers can be produced by a number of different techniques such as drawing from a melt, template

synthesis, phase separation, self-assembly and electrospinning (Teo and Ramakrishna, 2006;

Schnell et al., 2007; Kim et al., 2008). A major challenge consists in filling the conduit with the

orientated fibers.

One strategy consists of stacking 2D films of parallel fibers above one another. In one previous

study, this was done with 10 – 12 films of electrospun poly(acrylonitrile-comethylacrylate) fibers

inside longitudinally split polysulfone tubes, which were then sealed and implanted to bridge

17 mm sciatic nerve gaps in rats (Kim et al., 2008). More recently, an improved method for

manufacturing 3D fiber scaffolds was published (Yang et al., 2011). Highly aligned parallel fibers

were electrospun onto individual non-conductive frames using a conductive wire as target. The

holding frames with suspended microfibers were then placed across the surface of a hydrogel,

covered with a spacer frame and another layer of hydrogel was added. This process was repeated

several times to create a 3D microfiber scaffold (Yang et al., 2011). The procedure explained in

the present study is based on the 3D accumulation of fibers already during the spinning process.

This could be used also for successive stacking of fiber and gel layers (Figure 11d,e) or for

casting a gel over the 3D fiber array (Figure 11f).

While I modified the physical design of the target electrode to achieve a 3D fiber configuration, a

similar technique of “air gap electrospinning” also used two horizontally separated targets, each

connected to a common voltage with respect to the spinneret (Jha et al., 2011). To manufacture an

Chapter II

56

implantable scaffold, an exterior coating of poly(glycolic acid)/poly(lactic acid) copolymer was in

this case deposited onto the outside of the cylindrical fiber configuration. For this, a circular steel

plate was placed behind the fibers, and the spinneret was used for subsequent electrospraying of

the copolymer (Jha et al., 2011). A different technique was developed by Schlosshauer’s group,

who produced oriented PCL filaments by hot melt extrusion through nozzles with six-leaf cross

sections. After extrusion, the filaments were drawn in a separate process. This resulted in yarns of

several hundred filaments, each with six longitudinal grooves, which could be inserted in a

tubular conduit (Ribeiro-Resende et al., 2009). These structured filaments have been

functionalized in various ways including siRNA (Hoffmann et al., 2011) and were successfully

used to fill epineurial tubes in vivo (Hajosch et al., 2010). Using the method of injection molding

chitosan conduits were produced that contained several hundreds of longitudinally aligned PLGA

fibers (Wang et al., 2005; Ding et al., 2010).

The second approach to achieve internally structured implants aims at scaffolds with longitudinal

pores or channels. One method to achieve this is the application of strong magnetic fields which

orient collagen or fibrin matrices (Ceballos et al., 1999; Dubey et al., 1999; Verdú et al., 2002).

Another promising technique consists in the controlled freeze drying of collagen suspensions.

When the solvent is caused to freeze from one side of the solution, collagen is displaced, thus

resulting in collagen walls of longitudinally orientated guidance channels. The diameter of these

pores can be controlled to be in the range of 20 – 120 µm (Bozkurt et al., 2007, 2009; Möllers et

al., 2009). Structured collagen conduits, which are commercially available (Perimaix, Matricel

GmbH), were implanted as sciatic nerve grafts in rats, and axonal regeneration followed the

longitudinal pores of the artificial nerve guide. In these experiments better outcomes were

achieved by co-implanting Schwann cells (Bozkurt et al., 2012), though fiber containing implants

appear to achieve similar or better results without the need to implant live cells (Wang et al.,

2005; Hu et al., 2013).

4.2 Biological functionalization of the growth substrate

Although pure polymer fibers, such as PCL or poly(lactic acid), already proved useful as guidance

structures, several studies showed that functionalization either by covalent binding of ECM

peptides (Bockelmann et al., 2011) or blending with ECM proteins (Schnell et al., 2007)

improved the ability of such fibers to guide cell migration and axonal growth. In the present

experiments the surrounding gel already consisted of collagen with signal sequences that can

activate cellular integrin receptors. With the use of hydrogels that contain signaling molecules one

might therefore use artificial polymer fibers without the further need of functionalization. I found

that axonal growth along non-functionalized PCL fibers in a 3D collagen environment was

initially faster than in pure collagen, however, with time, that is after 1 week, this difference

disappeared. Possibly, activation of axonal integrin receptors by the collagen supported axonal

5 Conclusion

57

elongation whereas no such signaling was initiated by the PCL. This outcome suggests that

additional modification of the fibers with signaling molecules may be beneficial.

Coating of artificial polymer fibers with laminin or fibronectin enabled axons to grow faster than

the migration of Schwann cells (Hoffman-Kim et al., 2010). An easy technique to achieve such

functionalization consists of blending a synthetic polymer with fibronectin, collagen or laminin

when preparing the electrospinning solution. Cell culture investigations demonstrated this

approach to be even superior, with respect to certain aspects, to more sophisticated chemical

modification (Koh et al., 2008). Recently, an innovative method for the synthesis of covalently

functionalized poly(D,L-lactide-coglycolide) fibers has been published. The polyester was made

hydrophilic by adding star-shaped poly(ethylene oxide) followed by covalent binding of cell-

adhesion mediating peptides. The resulting fiber matrix was bioactive and promoted cell adhesion

via recognition of the immobilized binding motifs (Grafahrend et al., 2011). For functionalization

through blending or covalent modification in the electrospinning solution the present 3D spinning

setup with staircase-like targets can be used without further modifications.

When, as here, electrospun fibers are embedded in a biocompatible hydrogel, the preferential

adherence of axons to the electrospun fibers is an issue. In previous 2D studies aligned fibers

were placed on non-treated tissue culture plastic (Kim et al., 2008) or on cell repellent substrates

(Schnell et al., 2007; Klinkhammer et al., 2009), which left the cells little choice but to attach to

the fiber substrate. In 3D composites that consist of guidance fibers suspended in a supportive

matrix a suitable balance of cell surface attachment between the different material components

must be struck. With the imbedding of pure PCL fibers in a collagen gel the guidance effect or the

morphological alignment was maintained. It remains to be seen whether this remains when

additional growth promoting factors, for example neurotrophins, are introduced into the matrix

(Dalton and Mey, 2009). Different combinations of gels and modification of artificial polymers

can now be tested.

5 Conclusion

I report a technique for the 3D electrospinning of parallel microfibers that were suspended in

collagen gels and incorporated in a tubular nerve implant. In contrast to conventional hollow

nerve conduits, this nerve guide with aligned arrays of anisotropic fibers directed axonal

regeneration and Schwann cell migration in the longitudinal axis of the scaffold. The comparative

advantage of this design in comparison with alternative designs will now be tested in vivo.

1 Introduction

59

Chapter III

Assembly of artificial nerve implants and assessment of their

suitability to support nerve regeneration in vivo

1 Introduction

Traumatic injuries are frequently accompanied by damage to a peripheral nerve (PN). A complete

nerve transection considerably prolongs the dysfunction of an affected body part far beyond the

damage to bones or muscular tissue and hence impairs the patient’s autonomy, even when most of

the injured tissue has already been restored. As a result, it can aggravate the patient’s emotional

distress and produce relevant social and economic costs (Rosberg et al., 2005). Despite these

serious consequences, options for clinical treatment of PN injuries are limited (Lundborg, 2000;

Deumens et al., 2010). Since an end-to-end suture is only possible when the two stumps can be

brought together in a tensionless manner, the first choice for treating a PN gap is the insertion of

an autologous nerve graft (Meek and Coert, 2014). However, this approach implicates donor site

morbidity and may cause other complications associated with the necessity of a second operation.

For short nerve gaps it is possible to bridge the two stumps with saline-filled but otherwise empty

tubes of biodegradable materials. These are the only artificial nerve bridges that are clinically

approved at the moment (Meek and Coert, 2014). Their use circumvents the drawbacks of

autologous nerve transplants, however at the cost of a generally reduced prospect of regeneration,

and the disadvantage increases with the size of the gap (Isaacs, 2010; Konofaos and Ver Halen,

2013). Intensive research is therefore being carried out to develop new scaffold designs and

materials that allow functional recovery after PN injury (Mey et al., 2012; Jager et al., 2014).

The reduced efficacy of conventional hollow nerve conduits is attributed to the lack of supportive

cues, which are present in the distal nerve stump or in an autologous nerve graft (Meek and Coert,

2014). These include endogenous trophic factors acting on regenerating axons, including

neurotrophins, glial cell-derived neurotrophic factor (GDNF) or the fibroblast growth factors

(FGFs; Dalton and Mey, 2009). They can be incorporated with microspheres that gradually

release growth factors (Kokai et al. 2010). Alternatively, the delivery can be achieved by

introducing cells with the implant. This approach was recently chosen by Allodi et al. (2014), who

used a collagen matrix seeded with transfected Schwann cells overexpressing FGF-2 inside a

silicone tube to repair a 6 mm sciatic nerve gap in rats. Other cell types that support PN

regeneration are several types of stem cells (di Summa et al., 2010; Hu et al., 2013) and olfactory

ensheathing cells (Dombrowski et al., 2006; You et al., 2011). In addition, structural cues in the

extracellular matrix, most importantly the Schwann cell-derived basal lamina, play an important

role in axonal guidance (Johnson et al., 2011; Brooks et al., 2012). To provide a physical substrate

a number of researchers used biological materials, e.g. silk fibroin and collagen, or synthetic

Chapter III

60

polyesters, e.g. poly(ε-caprolactone) (PCL) and poly-glycolic acid (PGA). Bozkurt et al. (2007)

developed a collagen matrix with longitudinal pores to facilitate nerve regeneration, and used it to

reconstruct a biopsy induced sural nerve gap (Bozkurt et al., 2014). Particular attention has been

paid to the application of longitudinal filaments either as a part of the conduit’s wall or in its

lumen (Kim et al., 2008; Madduri et al., 2010b; Jha et al., 2011). In a clinical case study, Fan et

al. (2008) used a chitosan tube with internal filaments of PGA for repair of a 35 mm median nerve

defect in a patient. Despite these promising studies, the available artificial conduits are not

considered satisfactory, especially in comparison with autologous nerve grafts. Implants that

contain live cells are more successful, however, here the issues of immune suppression or of

finding an appropriate source for the cellular implants arise (e.g. Mey et al., 2012). For this

reason, I am aiming at a cell-free, artificial implant design that mimics structural features of the

natural ECM. In preceding experiments, I showed that arrays of electrospun PCL fibers aligned in

parallel determined the direction of Schwann cell migration and axonal growth. This effect could

be enhanced when fibers included biological signals, for example by blending PCL with collagen

(Schnell et al., 2007) or by covalent binding of integrin-activating peptides (Bockelmann et al.,

2011). This was shown first on two-dimensional substrates in vitro and recently transferred to

three-dimensional (3D) substrates with fibers embedded in collagen hydrogels (see chapter II or

Kriebel et al., 2014). The design developed in this study showed the potential of such fibers to

partly mimic basal lamina functions. Based on these results I designed an artificial conduit for PN

regeneration with an internal structure of sub-micron scale PCL or c/PCL fibers embedded in a

gelatin matrix.

In the present study, the suitability of these implants was tested by repairing a 15 mm sciatic

nerve gap in rats immediately after transection. Functional recovery was assessed via weekly

behavioral experiments followed by electromyography and histological evaluation of nerves,

implants and peripheral muscles 12W after surgery.


2.1 Manufacture of artificial implants

To manufacture structured implants of sub-micron scale PCL-fibers in a 3D gelatin matrix

(implant design referred to as “G1”), 12 wt% PCL in chloroform/methanol (3:1) were electrospun

onto a stepped target. In this way, a 3D configuration of fibers (Figure 16a and b). For more

detailed information on the setup, spinning process and fibers see chapter II or Kriebel et al.,

2014. Subsequently, sterile 5% gelatin, dissolved in PBS, was applied onto an object slide to form

a 1.5 mm thick film into which the fiber array was submerged (Figure 16c). Using hot forceps the

fibers were detached from the target, and the gelatin was kept on ice for 15 minutes. Two

razorblades with a fixed distance of 1.5 mm were used to punch out stripes of the gelled films


61

with longitudinal fibers. The stripes were frozen and inserted into hollow collagen conduits to

form the artificial implant. Empty spaces between conduit wall and internal scaffold were

cautiously filled up with gelatin solution followed by immediate cooling. For the second design

(G2) blended c/PCL fibers were spun from a solution of 12 wt% collagen/PCL (1:3), dissolved in

1,1,1,3,3,3-hexafluoro-2-propanol. The rest of the procedure was carried out as for G1. The empty

collagen conduits were provided by a commercial manufacturer and, for comparison, were used

also in a third group of non-structured implants (ET).

2.2 Animals and surgical procedure

For the experiments 10-week old, female Lewis rats of ca. 200 g body were used weight. Animal

housing and treatment was carried out in accordance with SfN guidelines for animal

experimentation and with permission of the Landesamt für Natur, Umwelt und Verbraucherschutz

Nordrhein-Westfalen. There were five experimental groups with 15 rats per group. Of these, three

were sacrificed after 2W for histological assessment of the biocompatibility of the implants, seven

after 12W (electrophysiology, behavioral assays, histology) and five after 14 weeks (behavioral

assays, neuroanatomical tracing, tracing data not included in this study). For surgery, anesthesia

was induced and maintained via inhalation of an isoflurane/air mixture of 5 % and 3 %,

respectively, using the Univentor 400 (Univentor High Precision Instruments). The depth of the

anesthesia was checked by testing of pedal withdrawal and corneal reflexes. The surgical area was

shaved and disinfected. Animals were placed in a prone position onto a thermostatically regulated

heating pad to maintain a constant body temperature. Subsequently, the right sciatic nerve, from

the sciatic notch to the point of trifurcation into peroneal, sural and tibial nerve, was exposed via

Figure 16 Brief documentation of the fiber production and embedding in gelatin. (a) Stair case-like target electrodes for electrospinning were attached to a U-shaped frame to allow handling of the fiber array. (b) Electrospun fibers accumulated at the target electrodes during the spinning process. (c) Gelatin (ca. 40°C) was applied onto a glass cover slip and fibers were submerged into the gelatin and detached from the target electrodes. Subsequently, the gelatin fiber composite was chilled to 4°C. From this, composite stripes were produced and inserted into collagen tubes to form implants G1 and G2.

Chapter III

62

incision of the skin and blunt dissection of the thigh musculature. This was followed by the

excision of a 15 mm piece of the sciatic nerve. According to experimental group, animals

immediately received one of the following treatments. For an autologous nerve transplantation

(AT) the excised piece of nerve was re-implanted with epineural sutures using 10-0 monofilament

nylon (Ethilon, Ethicon Inc., Somerville, USA). For groups G1, G2 and ET the respective

artificial implants were trimmed to a length of 17 mm and equilibrated for three minutes in cold

saline. Then the nerve stumps were inserted 1 mm into the ends of the implant and secured with a

single 10-0 mattress suture. Any residual air bubbles inside the conduits were removed before

insertion of the proximal stump. In the negative control group (lesion only, LO), the proximal

stump was embedded in the neighboring thigh muscle and secured with a 10-0 stitch to prevent

regeneration of axons into the distal stump. Finally, the thigh muscles and the skin were closed

with 7-0 and a 4-0 suture respectively. For pain relief animals received a subcutaneous injection

of 0.05 mg/kg body weight buprenorphine hydrochloride (Temgesic).

2.3 Tissue preparation and immunohistochemistry

At 2W after surgery and subsequent to electromyography at 12W after surgery, anesthetized rats

were transcardially perfused using 4 % buffered paraformaldehyde (PFA). Immediately after

perfusion, the sciatic nerve and muscles of the lower hind limb from both sides were dissected

post-fixed 24 h with 4 % buffered PFA and stored in PBS containing 0.05 % sodium azide.

Muscular atrophy: Since the susceptibility to atrophy depends on the predominant muscle fiber

type (Yamakuchi et al., 2000; Desaphy et al., 2001) four different muscles were weighed, which

vary in fiber type composition or size. The small Musculus soleus and M. extensor digitorum

longus have comparable size, but vary in fiber type composition with approximately 95 % slow

twitch/5 % fast twitch in the first and mostly fast twitch fibers in the second muscle (Soukup et

al., 2002). The M. gastrocnemius and its antagonist the M. tibialis anterior were more

heterogeneous and much larger. For muscle weight analysis, muscles were briefly dabbed dry

with tissue paper and weighed. Muscle weight on the operated side was normalized to the

contralateral muscle of the individual animal and subsequently averaged per group.

Nerve regeneration: For sectioning, nerve samples were cryoprotected with 30 % buffered

sucrose, embedded in Tissue-Tek and snap frozen in isopentane (-80 °C). Using a cryostat, cross

sections (16 µm) were made from the nerve proximal and distal to the implant (including tibial,

sural and peroneal branches) as well as from the medial part of the implant, and longitudinal

sections (16 µm) from the proximal and distal coaptation sites. For immunohistochemistry, the

sections were blocked for 1 h with 0.1 M PBS containing 5 % normal goat serum (NGS) and 0.5

% Triton X100, incubated overnight at room temperature in a mixture of primary antibodies,

diluted in 1 % NGS and 0.5 % TritonX100. A mouse monoclonal antibody was used against β-

tubulin III (T8660, Sigma, 1:1000), rabbit polyclonal antisera against S100 (S2644, Sigma,


63

1:1000) and choline acetyltransferase (ChAT; AB143, Millipore, 1:200). Subsequently, sections

were washed and incubated for 2 h with Alexa Fluor-conjugated secondary antibodies goat anti

rabbit IgG (Alexa Fluor 488, A11008, Invitrogen, 1:1000) and goat anti mouse IgG (Alexa Fluor

546, A11003, Invitrogen, 1:1000) and 4',6-diamidino-2-phenylindole (DAPI, Sigma-Aldrich; 1

µg/ml) in 1 % NGS and 0.5 % TritonX100. Histological sections were visualized with a Leica

TCS SP2 confocal microscope (Wetzlar, Germany).

2.4 Image analysis

Cross sectional areas of the regenerated tissue in the medial segment of the implant/autologous

nerve graft were determined using ImageJ 1.46r (NIH, USA). Briefly, the regenerated tissue was

encircled using the polygon selection tool followed by an area measurement of the region of

interest. One image per animal and all animals (N ≥ 5) per group were analyzed.

Axon numbers were counted and the cross sectional area of individual axons determined with

custom-made MATLAB software (TheMathWorks, Natick, MA). Images of cross sections of the

medial segment, stained against βIII-tubulin, were converted into binary images using individual

thresholds set by an observer blind to the treatment conditions of the samples. Thresholds were

chosen so that during conversion the original data of axon size were conserved as far as possible

while avoiding confluence of images of individual axons. From the converted binary images, the

software automatically determined number and size of individual axons. The obtained data were

checked with manual counting (in 2-3 samples per experimental group). To exclude profiles of

non-plausible sizes, the software was set to detect only objects within a range of 0.2 -78 µm2

(≈0.5 - ≈10 µm diameter).

For analyzing the percentage of motoraxons in the total axonal population, images of medial cross

sections from three individual animals per group, co-stained against βIII-tubulin and ChAT, were

used. The percentage of ChAT-positive axon profiles that were also positive for βIII-tubulin was

determined in randomly chosen regions of interest.

2.5 Electromyography

The experiments were conducted 12W after initial surgery. Anesthesia was performed as

described above, and body temperature was controlled using a heating pad. After exposing the

sciatic nerves on both sides, a bipolar hook electrode was used to stimulate the sciatic nerve were

proximal to the lesion, close to the sciatic notch. Compound muscle action potentials (CMAPs)

were recorded from the plantar Musculus flexor digitorum brevis and the M. interosseus pedis

dorsalis between toes two and three using a monopolar needle electrode. The reference electrode

was positioned at the contralateral site between the separated thigh musculature. Stimuli were

rectangular current pulses of 100 µs duration and 30 % supramaximal intensity, determined by

stepwise increase of the injected current until CMAP amplitude reached a plateau. Absolute

Chapter III

64

intensity was on average 1.76 ± 1.78 mA (mean ± SD). For a single measurement, ten repetitions

with a frequency of 1 Hz were averaged.

Signal detection and analysis was performed using custom made MATLAB software. For

detection of the evoked potential, a threshold slope and, in case of small signals, an amplitude

threshold was used. The software automatically analyzed the time to first peak, in the following

referred to as “latency”, and the potential difference between highest and lowest peak, referred to

as “amplitude”. To correct for differences in amplitude caused by a different impedance of the

recording electrodes, CMAPs on the operated side were normalized to the corresponding

contralateral side of the same animal, which was recorded immediately afterwards. For the final

evaluation, means and SD (N ≥ 5 animals) were calculated for every experimental group.

2.6 Behavioral experiments

To assess the recovery of motor functions, toe spread and grip strength baseline values were

determined on three consecutive days prior to surgery, followed by weekly testing starting 1W

after surgery.

For toe spread measurements, animals were placed on a transparent surface with a camera system

underneath. The animals were lifted up by the experimenter and a sequence of images with 1 Hz

frequency was produced to gather at least six images, in which the animal was not moving and

had both feet in one plane. Custom made MATLAB software was used to present the images in a

randomized fashion to a single blinded observer, who had to mark the position of the 1st and 5th

toe of both feet. Subsequently, the software determined the mean ratio of the distances between

the toes of the operated and non-operated hind limb.

Grip strength was determined using the “Bioseb grip strength test” device. For this, while the

animal was held by the experimenter, its hind paw was placed on a metal grid fixed to the test

device. Then the animal was pulled away from the grid, while the maximum force which the

animal could exert on the bars was measured. Three trials were averaged per animal and time

point. For the final analysis of toe spread and grip strength measurements, means and SD of all

animals (N ≥ 11) in the respective groups at a given time point were calculated.

2.7 Statistical analysis

Individual animals were considered as independent samples, and means, SD and SEM were

determined for the experimental groups and time points. Data processing, except for

electromyography, and statistical analysis were performed using IBM SPSS Statistics Version 21

(IBM Corp., Armonk, NY, USA). Intergroup differences were tested with ANOVA. After

Levene’s test, post-hoc comparisons were performed according to Dunnet T3 for unequal

variances (i.e. analysis of grip strength, toe spread, muscle weight and median axon size) and

Scheffé for equal variances (i.e. analysis of nerve cross section area and axon count). Statistical

3 Results

65

analysis of electromyography data was performed with the MATLAB software. Intergroup

differences were tested with ANOVA using post-hoc Bonferroni-corrected t-tests.

3 Results

3.1 Artificial nerve implants and general observations

Based on a recently published technique (Kriebel et al., 2014) artificial nerve guides were

produced that consisted of a biodegradable tube with a three-dimensional array of longitudinally

oriented PCL and c/PCL fibers embedded in a gelatin matrix (Figure 16).

Following surgery the rats recovered well and did not display signs of pain, self-mutilation or

secondary complications, such as inflammation, throughout the study. Three animals of each

experimental group were sacrificed at

2W after implantation. In all cases

the nerve guides or autologous

transplants appeared well integrated

within the host tissue and showed no

macroscopic signs of neuroma

formation, or excessive encapsulation

of the conduit by fibrotic scar tissue.

At this point in time (2W), the

collagen tubes were still intact, but light-microscopical investigation revealed no vestiges of the

electrospun fibers. At 12W after surgery the tubes had been completely resorbed in all animals,

whereas the non-degradable suture material was still visible (Figure 17).

Figure 17 Tissue regeneration following sciatic nerve transection (15 mm gap size) and subsequent repair using different nerve bridges. The photographs show (a) a non-lesioned sciatic nerve (contralateral control, CL),b) regenerated tissue 2 weeks after the insertion of an artificial implant (as an example for all artificial conduits, which looked similar) and (c) regenerated tissue 12 weeks after the insertion of an autologous nerve graft (AT) and different artificial implants (ET, G1 and G2; see text).

Chapter III

66

3.2 Tissue regeneration: macroscopic evaluation of the implants

Gross examination of the artificial nerve conduits (groups G1, G2 or ET) showed that the conduit

was surrounded by a thin layer of connective tissue. At the proximal suture sites, the nerve stumps

were slightly swollen, visibly sealing the ends of the tubes. Inside the conduit a tissue cable with

clearly visible blood vessels had formed spanning the entire length of the lesion (Figure 17b). At

the later time point (12W), when the collagen tube had been degraded, the swelling at the suture

site was reduced at the proximal and had completely disappeared at the distal interface. At both

proximal and distal coaptation points, a network of blood vessels emerged. The regenerated tissue

between the proximal and distal nerve stumps had increased in diameter compared to the earlier

time point in all experimental groups (except the non-implanted group, which lacked a

regenerated tissue connection between the stumps; not shown). After 12W there was no apparent

difference between implants G1, G2 and ET, which were still thinner than the autologous nerve

grafts. Macroscopically, the autografts resembled the non-operated sciatic nerves (Figure 17c).

3.3 Schwann cell infiltration and axonal regeneration in the implants

To assess the regeneration inside the implants, longitudinal sections of the coaptation sites

between nerve stumps and the implants, and cross sections of the proximal, the distal nerve

stumps and of the mid-section of the implanted conduits were produced.

An analysis of the cross sectional area of the different implants after 12W confirmed the

macroscopic assessment of substantial growth in all artificial nerve bridges (Figure 18). The

regenerated tissue was similar in size in all experimental scaffolds (ET, G1, and G2) and still

significantly smaller than the resected and

re-implanted nerves (AT), which had the

same diameters as the non-operated control

nerves.

The inspection of the proximal connections

between nerves and implants 2W after

surgery indicated a better ingrowth of

axons (III-tubulin labeling) and Schwann

cells (S100 staining) into the internally

structured implants (G1, G2) than into the

empty tubes (ET; Figure 19). All groups

had intense staining of axons and S100+

cells proximal to the lesion. Substantial

infiltration of Schwann cells and axonal

elongation was visible in G1, G2 and AT

grafts at the proximal interface, and some

Figure 18 Cross sectional area of the regenerated tissue in the medial segment 12W after nerve transection and repair. The cross sectional areas of the autologous nerve grafts (AT) were unchanged 12W after nerve repair, resembling the non-operated situation (CL), tissue that regenerated through one of the artificial implants had significantly smaller cross sections. There were no differences between ET, G1 and G2. Symbols indicate significant differences against groups CL (*) and AT (#) (p** < 0.01; p*** < 0.001; N ≥ 6).

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axonal and S100+

profiles could be found in cross sections of the medial segment. In the ET

group, in contrast, the staining intensity declined abruptly at the proximal interface. At 2W there

were few axons and S100+ cells penetrating the ET conduit, but none were detected in the

midsection of the implanted segments. Distal to the graft no axons were found in any of the

groups at this early time point (2W; Figure 19).

The same assessment was performed 12W after surgery (Figure 20). In the physiological

situation, represented by the contralateral, non-lesioned sciatic nerve (CL), axons were evenly

distributed in the nerve and grouped into a few fascicles. The autograft after 12W showed a

higher degree of fasciculation but generally resembled the non-lesioned nerve. In the artificial

Figure 19 Schwann cell infiltration and axonal regeneration at the proximal interface between nerve and implant 2W post surgery. Longitudinal sections from groups CL (top), ET (middle) and G2 (bottom) stained against βIII-tubulin (a, c, e) and S100 (b, d, f) to visualize axons and Schwann cells, respectively. Yellow arrowheads point to the collagen tube, which is still present 2W post surgery. The weaker staining intensity beyond the interface of nerve and implant in group ET compared to G2 indicates a slower infiltration of Schwann cells and axons into the implant.

Chapter III

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implants multiple small fascicles were loosely distributed over the entire cross sectional area, each

of them enveloped by a thin layer of connective tissue (Figure 20g-l). In the center of eight of the

internally structured artificial implants a longitudinal compartment was found, which was devoid

of axons and Schwann cells but filled with S100III-tubulin

cells (data not shown). Although

this structure could occupy an area of up to 25% of the regenerated nerves, there was no

correlation with the quality of functional regeneration.

Quantification of the number of III-tubulin+ profiles in cross sections of the medial segment

Figure 20 Schwann cell infiltration and axonal regeneration in the medial segment of the regenerated nerve 12W post surgery. Cross sections of the medial segment from groups CL, AT, ET and G2 (from top to bottom) were stained against βIII-tubulin (a, d, g, j) and S100 (b, e, h, k). Compared to CL and AT, nerves that regenerated in the artificial implants show a smaller cross sectional area and increased fasciculation. Images of the co-staining with both antibodies at higher magnification are shown in (c), (f), (i) and (l), with the biggest axons in AT (apart from CL) and the smallest in ET. Schwann cells enwrapping single axons can be seen in AT and G2 but less so in ET.

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revealed no differences between the non-injured nerves and the autologous implants, whereas the

artificial implants contained significantly fewer axons (Figure 21a). In all three groups (ET, G1,

G2) about 3.400 axonal profiles per cross section were counted, about half of the number

contained in the control nerves. Interestingly, the number of III-tubulin+ profiles in the medial

segments of the AT group was higher than in control nerves (CL) and also higher than proximal

to the lesion, even though this difference was not statistically significant. In all nerve repair

groups with artificial implants an increase in the number of axonal profiles was observed from the

midsection of the implant to the distal segment (significant for G1 and G2; p < 0.05 and p <

0.001, respectively), suggesting that sprouting of regenerating axons at the distal coaptation point

had taken place.

Apart from the number of all regenerating axons, the efficacy of the different implants to support

regeneration by motor neurons was investigated. Using immunostaining against ChAT (Figure

22a-b) the percentage of motor axons among the total population of axons (III-tubulin+, i.e.

motor, sensory and autonomic axons) in the midsection of control and regenerated nerves were

Figure 21 Quantitative analysis of the regenerated nerve tissue. Panel (a) shows the number of individual βIII-tubulin+ profiles (regenerated axons) in cross sections of the medial segment (significant difference against CL and AT are indicated as * and #, respectively; p<0,001; N≥6), (b) size distributions of βIII-tubulin+ profiles, step size 1µm2, and (c) their cross sectional areas in µm2 in the different groups. The solid and dashed lines in c indicate median and mean, respectively; the boxes 25 - 75 %, whiskers 10 – 90 % and dots 1 % and 99 %.

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70

determinded. Depending on the experimental group this percentage ranged from 32.1 ± 1.2 %

(G1) to 37,1 ± 5.4% (CL) with no significant difference between any of the groups or the non-

operated control nerve (Figure 22c).

3.4 Size distribution of regenerated fibers

High magnification images revealed considerable differences in the thickness of axons and

association of Schwann cells with the fibers in the different groups. Cross sections of non-

lesioned sciatic nerves showed many large and roughly circular or elliptical axonal profiles that

were regularly encircled by S100+ Schwann cells (Figure 20c). Non-lesioned nerves exhibited

regions of predominantly bigger or smaller axons with a gradual transition between them. This

was different in the tissue 12W after nerve repair (Figure 20f, i, l) Axons were generally smaller

and the fewer large diameter axons were randomly intermingled with the mass of small axons.

The diameters of III-tubulin+ profiles were for CL: 6.0 ± 3.2 µm (mean +/- SD), AT: 3.4 ± 2.0

µm, G1: 3.9 ± 1.8 µm, G2: 5.1 ± 2.1 µm, and ET: 3.1 ± 1.5 µm. The size distribution of III-

Figure 22 Similar percentage of regenerating motoraxons in all implants. Immunostaining against βIII-tubulin (red) and ChAT (green) was used to identify motoraxons in cross sections from repaired (mid of lesion) and non-operated sciatic nerve samples (exemplary image from CL). Only motoraxons are characterized by co-localization of both antibodies (examples: white arrowheads). Quantification of the ratio of motoraxons to the total axonal population revealed no difference between the different experimental conditions and the non-operated contralateral nerve.

3 Results

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tubulin+ profiles (Figure 21b) confirmed that small axon diameters were relatively more abundant

in the regenerated nerves. When comparing these groups, I found the number of large caliber

axons (larger than 10 µm2 cross section) to decrease from AT (548 ± 94; mean ± SEM) to G2

(389 ± 44) and G1 (307 ± 74) to ET (209 ± 29). For the non-lesioned nerve the value was 1378 ±

130. The median axon diameter was largest in the non-injured nerves, and this was significantly

different only from the ET group (p < 0.05), which had the smallest fibers (Figure 21c). The only

other experimental group with a significantly larger median of axon diameters compared to ET

was the G2 construct. It was not the case for AT because of a large number of regenerated small

fibers.

As with the axons, the appearance of S100+ positive profiles was less regular in the repaired

situation compared to control nerves. However, most S100+ positive profiles in samples from

group AT displayed a clear circular shape suggesting a myelinating phenotype, and this was

confirmed with staining against myelin basic protein (data not shown). This pattern was also

observed in groups G1 and G2, however there were also Schwann cells that did not show a clear

interaction with one axon, thus likely to have a non-myelinating phenotype. This observation was

even more pronounced in group ET (Figure 20i). In summary, judging the maturity of regenerated

axons by diameter and state of myelination, I saw the highest similarity with the non-lesioned

nerve in the AT group, a lesser degree in the structured implants G1, G2, and very little of this

organization in the ET control group.

3.5 Functional regeneration: electromyography

At the endpoint of this study, 12W after sciatic nerve transection and repair, functional

regeneration was assessed by recording of CMAPs from the plantar Musculus flexor digitorum

brevis (Mfdb) and M. interosseus pedis dorsalis (Mipd), both innervated by the tibial branch of

the sciatic nerve. CMAPs were induced with electrical stimulation of the sciatic nerve proximal to

the lesion site (Figure 23a, traces of typical recordings). The successful recording of induced

electrical muscle activity in all rats that received autologous transplants (AT) or scaffolds with

PCL-fibers (G1) and in five of the six animals that received c/PCL-fiber constructs (G2)

demonstrated that the new artificial implants supported nerve regeneration including re-

innervation of target tissue. In contrast, I was only able to record CMAPs in half of the animals in

the group that received an empty collagen tube (ET). As anticipated, no CMAPs could be

recorded from the negative control group (lesion only, LO; Figure 23b).

To evaluate the functional quality of regeneration in more detail, I analyzed amplitude and latency

of the recorded potentials (Figure 23c, d). The CMAP amplitude is dependent on the number of

functional motor units and the temporal coherence of nerve excitation in the vicinity of the

electrode, hence gives an indication of the number and quality of regenerated motor axons. For

group AT I found the strongest recovery with 48 % of the CMAP amplitude recorded, compared

Chapter III

72

to the non-lesioned contralateral side (Mipd). The second highest CMAP amplitudes were found

in G2 (22 %) followed by G1 (16 %) and the lowest in ET (10 %, Table 1). The differences

between the artificial constructs were not significant.

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Figure 23 Electromyography performed 12W after sciatic nerve repair. (a) Exemplary recordings of the compound muscle action potential (CMAP) from Musculus flexor digitorum brevis (Mfdb) for different experimental groups after sequentially increasing stimulus intensities (grey lines) up to a 30% supra-maximal stimulus (black line). The arrow indicates the time of stimulation. (b) Percentage of animals that showed CMAPs indicating successful reinnervation. The number in white bars indicates the sample size while the colored bars indicate the number of animals with evoked CMAPs from at least one muscle (Mfdb or Mipd). (c) CMAP amplitude of Musculus interosseous pedis dorsalis (Mipd) on the lesioned side was normalized to the contralateral (non-lesioned) side (grey dots). (d) CMAP latency (time to first peak) of Mipd on the lesioned side and the corresponding contralateral side (grey dots). Black bars in c, d indicate mean ± SD; asterisks mark significant differences with respect to group AT (p** < 0.01; p*** < 0.001; N ≥ 5). The difference to the contralateral side within individual groups was always significant (p < 0.001). (e) Data from both muscles were pooled and CMAP amplitude plotted against latency. The graph reveals a clear negative correlation between both parameters, and recordings from individual groups are distinctly clustered.

Chapter III

74

Table 1 Compound muscle action potentials after PN repair with artificial implants Data are means ± standard deviation. The number of rats of whom CMAPs could be recorded were for AT: n = 6, ET: n = 3 G1: n = 7, G2: n = 5 (animals without CMAP recording in groups ET and G2 are excluded from latency data). CMAP amplitude of the non-lesioned nerve was set to 100 % individually.

The maturity of regenerated axons can be assessed by measuring CMAP latency, which are

considered to be influenced by axon diameters and the degree of myelination. The same relative

ranking between the experimental groups that I found with response amplitude was also observed

when comparing latencies between electrical stimulation and CMAP (Table 1). Shortest latencies

were determined in group AT. While differences between the three artificial implants were not

significant, the shorter mean latencies of construct G2 compared with G1 and ET again indicated

an advantage of the c/PCL-fiber containing matrix.

Latency and amplitude of the CMAPs were negatively correlated. The CMAPs of non-lesioned

nerves had relatively high amplitudes (many motor units were activated simultaneously)

combined with short latencies (high conduction velocities of the stimulated fibers), whereas a

combination of low amplitudes and long latencies were seen when regeneration was poor.

When plotting CMAP amplitude vs latency the graph describes a hyperbola that asymptotically

approaches the abscissa (Figure 23e). Data of the non-injured controls and the two implant groups

AT and G2, form clusters along this function, illustrating the rank order between these groups that

are characterized by successful recovery. Data points for groups ET and G1 tend to form a single

cluster at the bottom end of the function indicating an inferior recovery according to

electromyography, compared to the other treatments.

CMAP amplitude

[% of control] CMAP latency

[ms]

M. flex. dig. brev.

M. inteross. ped. dors.

M. flex. dig. brev. M. inteross. ped.

dors.

non lesioned side (CL)

100 100 3.21 ± 0.22 3.07 ± 0.31

lesion only (LO) 0 0 n.a. n.a.

autograft (AT) 41.4 ± 5.5 47.7 ± 10.3 4.67 ± 0.24 4.71 ± 0.26

empty tube (ET) 5.7 ± 8.9 10.1 ± 11.4 6.23 ± 1.03 6.11 ± 0.28

PCL-implant (G1) 17.1 ± 7.9 15.6 ± 6.9 6.65 ± 0.51 6.74 ± 0.65

C/PCL-implant (G2) 21.5 ± 15.8 22.1 ± 16.4 5.81 ± 1.01 5.89 ± 0.98

3 Results

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3.6 Functional regeneration: recovery of motor functions

The crucial outcome with respect to PN regeneration is the recovery of appropriate motor

functions. To investigate the implications of nerve transection and the progress of regeneration

after nerve repair using different implants, behavioral experiments were performed weekly

starting 1W after surgery. At that time point, all animals exhibited paralysis and avoided the use

of the lower hind limb on the operated side. Behavioral observations of the rats confirmed the

completeness of the nerve transection and the gradual regeneration. Two quantitative measures,

toe spread on response to lifting the animal and grip strength, revealed when the animals regained

motor functions of their hind paws.

The toe spread score measures the ability to extend and abduct the digits. In the days following

surgery the affected paw showed a flat or passive posture with extended toes and reduced spread

between digits 1 and 5 compared to baseline (Figure 24 a, b). In the following weeks, toe spread

decreased further, reaching a minimum of approximately 30 % of baseline 3W after surgery. For

the animals with sciatic nerve lesion but no implant, the toe spread value remained at this value

without improvement until the end of the experiment at 12W. In the animals with nerve implants,

toe spread began to increase and by 12W had recovered to a large extent (up to 75% compared to

the contralateral paw). Compared to the non-repair group this recovery was significant for groups

AT, G1 and G2 but not for the ET group (Figure 18b).

The grip strength test measures the muscular force exerted by the digits when flexing or grasping

a metal grid. Consistent with the paralysis in the operated lower hind limb, all animals failed to

exert any additional force beyond mere friction as the paw moved across the metal bar surface

during the first weeks after surgery (i.e. below 15 N, dashed line in Figure 24c). Starting at 6W

after surgery, the first animal (in group AT) exhibited a change from an extended to a flexed

posture of the affected toes. Mean grip strength of group AT started to recover 7W after surgery

and, from then onwards, showed a steady and linear increase. The groups with artificial nerve

grafts also showed regeneration starting to recover in approximately one week intervals, with

group G2 being first, followed by G1 and ET last. In these cases the observed functional recovery

was more modest, starting later and developing with a lower slope (Figure 24c). At the endpoint

(12W after surgery) mean grip strength for all groups with the exception of ET was significantly

larger than that of the group without implant (LO, Figure 24d). This outcome was reflected in the

higher number of ET individuals compared to the other groups, that did not exhibit grip strength

above 15 N, considered as no recovery, during the time of the study. While this threshold was

reached by 100 % (11/11) of the individuals in groups AT and G1, and by 75 % (9/12) in group

G2, only 25 % (4/12) in ET exhibited a detectable increase in grip strength.

Chapter III

76

Figure 24 Behavioral analysis of nerve regeneration. Following surgery the recovery of the toe spread reflex and grip strength was investigated once per week. (a) The toe spread reflex was investigated by measuring the distance between toes 1 and 5 on the operated and non-operated legs. (b) The toe spread ratio of the operated to the non-operated side is plotted for the different conditions at the 3W (white bars) and 12W (colored bars) time point. It drops immediately after the operation, declines further to a minimum at 3W for all groups and recovers in groups AT, G1 and G2. (c) Grip strength development over time illustrates the degree of functional regeneration. A threshold of resistance without muscular activity (dashed line) was never surpassed by any rat of group LO. Plot (d) shows the grip strength before surgery (baseline) as well as 1W (white bars) and 12W (colored bars) post-surgery, respectively. Error bars indicate SEM, asterisks significant differences to the mean of group LO 12W post surgery (p** < 0.01; p*** < 0.001; N ≥ 11).

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3.7 Muscular atrophy and regeneration

A secondary effect of the denervation after nerve transection is muscular atrophy due to the lack

of activity. To the extent that this is caused by the degradation of proteins in muscle fibers which

survive, reinnervation can initiate new growth of muscle mass. To include different muscle fiber

types in the analysis the weight of four muscles of the lower hind limb were measured, i.e. M.

soleus, M. extensor digitorum longus, M. gastrocnemius and its antagonist the M. tibialis anterior.

(Table 2 and Figure 25). Denervated muscles lost approximately 50 % of their weight in the first

2W after nerve transection. This was similar in all experimental groups and independent of the

type of implant the animals had received. With sustained denervation, as in the non-repaired

situation (LO), muscle weight decreased further, reaching 26 % (average of all four muscles) after

12W. In animals that received one of the artificial implants this continuing decline in muscle

weight was prevented. After termination of the experiment their muscles were only slightly

smaller than 2W after surgery, and this difference was not significant. Given that reuse of re-

innervated muscles began later than 6W even in the AT group (Figure 24c), these results indicate

a recovery of muscle weight after reaching a minimum sometime between 2W and 12W. Only

animals treated with an autologous nerve graft (AT) showed considerable recovery of muscle

weights beyond the values obtained 2W after surgery. The average muscle weight in group AT

reached 71 % of the non- lesioned side.

Figure 25 Determination of muscular atrophy. Following nerve transection the denervated muscles start to atrophy. (a) Tibialis anterior muscles from the operated leg 12W post surgery of experimental groups (LO, AT, G1) exhibit different degrees of atrophy. For comparison, the corresponding contralateral muscle (CL) is shown. (b) Atrophy was quantified by determination of the weight of four muscles of the lower hind limb (see text) at 2W (white bars) and 12W (colored bars) post surgery. Muscle weight of the operated leg was normalized individually to the contralateral side. For the graph the relative weight of all four muscles of every animal was averaged. Error bars indicate SEM, asterisks depict significant difference to LO at the respective time point (p*** < 0.001; N2W ≥ 2; N12W ≥ 6).

Chapter III

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4 Discussion

The quality of autologous nerve grafts to support functional recovery after PN injury remains

unmatched. Yet, their disadvantages, including the need for additional operations and the limited

availability of donor material create a demand for alternatives based on the use of acellular,

allogenic implants or artificial implants. In the present study, I tested the efficacy of an internally

structured scaffold to promote PN regeneration over a 15 mm sciatic nerve gap in rats. These

artificial implants contained aligned sub-micron scale PCL-fibers as longitudinal guidance cues

embedded in a gelatin matrix. It was derived from the design I documented in chapter I (see also

Kriebel et al., 2014), but with modifications regarding the gelatin hydrogel for embedding the

fibers (instead of collagen), and the choice of the surrounding tube. In its molecular composition

gelatin is closely related to collagen. I found that the gelling characteristics of gelatin facilitated

implant assembly and storage of the prepared scaffolds, while offering the possibility to sterilize it

before use. In addition, gelatin is much cheaper than collagen. For the hollow nerve conduit I used

a collagen tube instead of the previously used hand-made PCL tubes. Based on the recovery of

motor functions and on electrophysiological assessment of the regenerated nerves the structured

implants represent a significant improvement on the conventional hollow conduits and offer

possibilities for further optimization. A modification of the fibers by blending the PCL with

collagen before electrospinning produced an additional positive effect on regeneration.

Table 2 Muscle weight normalized to the corresponding muscle on the contralateral side 2 and 12 weeks after respective treatment. Data are % of individual contralateral muscle (CL) given as means ± standard deviation.

M. gastrocnemius

M. tibialis anterior

M. soleus M. extensor

digitorum longus

2

weeks 12

weeks 2

weeks 12

weeks 2

weeks 12

weeks 2

weeks 12

weeks

non lesioned side (CL)

1,38g ± 0,18g

1,16g ± 0,09g

0,51g ± 0,10g

0,40g ± 0,04g

0,13g ± 0,01g

0,12g ± 0,01g

0,12g ± 0,02g

0,09g ± 0,01g

lesion only (LO)

47,76 ± 5,20

17,30 ± 1,92

47,12 ± 3,59

24,24 ± 1,90

46,73 ± 4,08

34,69 ± 6,97

48,78 ± 6,22

29,13 ± 1,51

autograft (AT) 43,69 ±

1,43 55,95 ±

5,69 48,07 ±

0,06 72,94 ±

3,57 50,82 ±

1,39 69,50 ± 14,56

46,14 ± 0,16

86,06 ± 6,70

empty tube (ET)

44,80 ± 1,06

31,64 ± 12,06

50,46 ± 4,80

46,97 ± 19,79

51,11 ± 0,47

42,40 ± 19,92

51,63 ± 0,89

51,36 ± 18,90

PCL-implant (G1)

47,50 ± 2,60

40,65 ± 8,37

54,66 ± 4,37

49,33 ± 10,12

53,95 ± 7,35

45,48 ± 6,69

58,73 ± 0,76

60,62 ± 8,98

C/PCL-implant (G2)

48,45 ± 0,68

37,74 ± 9,91

53,76 ± 5,49

50,45 ± 16,63

51,30 ± 3,91

43,18 ± 11,60

67,30 ± 3,74

57,81 ± 18,53

4 Discussion

79

4.1 Regeneration in the artificial nerve implants

Following a PN crush lesion, the distal sections of the transected axons degenerate, and activated

Schwann cells and macrophages phagocytose cellular debris (Deumens et al., 2010). Structural

reorganization follows, characterized by a longitudinal alignment of proliferating Schwann cells

and basement membranes. These form bands of Büngner, which guide regenerating axons from

the proximal stump to targets in the periphery (Christie and Zochodne, 2013). Similar processes

promote regeneration in an autologous nerve transplant. Although the ECM and secreted factors

of an injured PN create a permissive environment for axonal growth, regeneration is initially

slow, because proximal axon endings degenerate back to the last node of Ranvier before the

metabolic changes take place that induce the formation of growth cones (Deumens et al., 2010).

Even in case of a direct end-to-end suture after sciatic nerve transection in mice it was reported

that only 17% of the axons had crossed the repair site after five days (Witzel et al., 2005).

When a large nerve gap is bridged by fluid-filled tubular implants, bands of Büngner are not

available, which complicates regeneration: Within the first week after surgery a reconnection of

the proximal and distal nerve stumps is initiated by the formation of a cord of longitudinal fibrin

(Weis et al., 1994; McDonald and Zochodne, 2003). How this first tissue bridge is produced in

detail remains to be elucidated. Several publications indicate that perineurial glia cells and

fibroblasts are involved in the process (Scaravilli, 1984; Weis et al., 1994), and my observation of

central strands filled with S100-/IIItubulin

-/MBP

- cells in regenerated nerves within artificial

conduits at 12W may reflect it.

Subsequently, Schwann cells infiltrate the lumen of the conduit. Migrating along the bridge of

connective tissue, they proliferate and deposit other ECM material including a laminin-containing

basement membrane. On this substrate regenerating axons from the proximal stump follow up

closely (Zochodne, 2012). In my experiments, axons were detectable at 2W in the medial segment

of the internally structured implants, but they were few compared to the endpoint observation at

12W. This may be expected, given that previous studies (Witzel et al., 2005) demonstrated that

axonal regeneration does not occur simultaneously but successively, with a few pioneering axons,

which are followed by consecutive waves of regenerating fibers. Based on the histological

evidence of an internal connective tissue fascicle after 12W that was surrounded by Schwann cells

and axons, I surmise that the formation of a connecting fibrin cable occurred also in the gel- and

fiber-containing implants. Although the physiological and functional data demonstrated better

regeneration in the structured implants compared to the merely fluid-filled tubes, the histological

evaluation was not suitable to demonstrate that axons or Schwann cells actually moved in contact

with these substrates. It is also possible, that the positive effect of the internal structure was

primarily based on aiding the formation and stabilization of the fibrin cable, the very first step in

regeneration, as suggested by Koh et al. (2010). Axon counting revealed that not all axons entered

into the artificial implants but that axonal branching must have occurred when axons crossed the

Chapter III

80

distal coaptation site and regrew into the denervated distal stump. Branching of fibers was also

observed within the autologous transplants at the proximal coaptation site, as reported before

(Jenq and Coggeshall, 1985; Miyamoto et al., 1985; Witzel et al., 2005). This observation

suggests that branching of axons is induced through contact or trophic factors released from a

nerve undergoing Wallerian degeneration, which could be either the distal stump or an autologous

nerve graft.

Since the time needed for reinnervation of a target tissue depends on the type, -size and location

of injury, it is difficult to compare the speed of recovery in the present implants with data from

the literature (Cao et al., 2011; Yu et al., 2011; van Neerven et al., 2012; Kim et al., 2014).

Nevertheless, my analyses of grip strength recovery and muscular atrophy are in accordance with

previous studies. After making a 10 mm sciatic nerve lesion in rats at mid-thigh level and

subsequent repair with an autologous nerve graft, a duration of 6-7 weeks until reinnervation of

the gastrocnemius muscle has been reported (Korte et al., 2011; Kim et al., 2014). Other

researchers (McCallister et al., 2001) autografted a 10 mm gap in the rat sciatic nerve and

observed a remaining muscle weight of the M. tibialis anterior of 62 % at 12w after injury (73%

in the present study) while the controls without repair had 29 % (24 % in the present study). The

better results of the present autograft is likely to be due to the fact that McCallister and colleagues

reversed the orientation of their short nerve graft before re-implantation, whereas larger lesions of

the present study and also those of Bozkurt et al. (2012) made the reversal of the graft impractical,

and therefore was not performed. Given the general agreement with published data concerning

autologous transplants the data about the structured artificial implants presented here may be put

into perspective. Thus, the novel PCL- and c/PCL-fiber containing gels represent a significant

improvement on standard implants without internal guidance structure, although they do not reach

the quality of autologous nerve grafts or allografts (Brooks et al., 2012).

4.2 The importance of an internal guidance structure in artificial implants

The only FDA and CE approved alternatives to autologous transplants are simple tubes

applicable for short distance repair (less than 2-3 cm). As corroborated here, they are inferior to

autologous transplants, and the integration of molecular and structural cues for reaching a

comparable functional outcome remains a challenge (Lundborg, 2000; Isaacs, 2010; Konofaos

and Ver Halen, 2013). Since fluid-filled tubes provide no substrate for axonal growth in their

lumen, it is only after some time, when a continuous tissue bridge has formed and infiltration of

perineurial and Schwann cells has occurred, that axons can start to grow into the implant (Lewis

and Kucenas, 2014). On the other hand, the potential for successful regeneration diminishes with

the time after injury (Xu et al., 2014), implying that the presence of a growth substrate may be an

important factor for functional recovery within a critical window of opportunity. I therefore

attribute the better functional outcome in my experimental groups to the incorporation of

4 Discussion

81

electrospun fibers that serve as longitudinal guidance cues (as shown in vitro: Schnell et al., 2007;

Kriebel et al., 2014) and not to the gelatin substrate surrounding them. This statement does not

diminish the importance of the choice of the hydrogel and its physical properties. Generally, cells

need surfaces to grow on and migrate, and in situations when the hydrogel forms a distinct

boundary with the surrounding medium, as observed with alginate and fibrin gels, cells tend to

grow along this interface rather than penetrating the gel (Mohanna et al., 2003; Gerardo-Nava et

al., 2014). This prevents regeneration within the intended structure. To avoid such effects I

decided to use non-fixed gelatin as a support for the fiber arrays. Kept cold, it allowed handling

and incorporation of the fibers into the tube, while at body temperature it became liquid again and

thus did not pose an obstacle for cells involved in regeneration.

4.3 Implementation of biological signals in synthetic cell-free nerve grafts

Under the condition of nerve crush with intact perineurial sheaths (axonotmesis) multiple

molecular signals on the ECM and molecules released from the Schwann cells induce axonal

regeneration. Hydrogels and electrospun fibers can provide a platform to present selected signals

in artificial implants (Dalton and Mey, 2009). Soluble growth factors, e.g. nerve growth factor

(Xu et al., 2003) and glial growth factor (Mohanna et al., 2003) were successfully used in this

context. Another option consists in functionalization of implanted surfaces with activators of

integrins and cell adhesion molecules (Klinkhammer et al., 2010; Grafahrend et al., 2011).

Already the implementation of one type of integrin-activating peptide (RGDS) on electrospun

fibers increased their ability to induce Schwann cell migration and axonal growth (Bockelmann et

al., 2011).

Considering that the implantation of living cells has serious practical disadvantages, such as

additional operations for harvesting endogenous cells or immune suppression, my strategy

focuses on the development of cell-free artificial nerve bridges. Two recently developed acellular

products that contain internal guidance structures are the Avance® and Neuromaix® nerve

implants. Both incorporate some biological activity because they are derived from natural

materials. The Avance (AxoGen Inc., Alachua, FL) nerve graft is marketed as the “only off-the-

shelf commercially available processed nerve allograft intended for the surgical repair of PN

discontinuities”. Recovered from human PN, which are cleaned from cellular debris, the grafts

preserve the ECM structure of the PN and have been reported to show similar axon regenerating-

promoting properties as autologous nerve grafts (Brooks et al., 2012). Neuromaix (Matricel

GmbH; Herzogenrath, Germany) consists of an orientated scaffold with longitudinal pores that is

produced from porcine collagen type I using a patented freezing protocol. Animal experiments

with this scaffold were also successful (Bozkurt et al., 2012), and a clinical phase 1 trial is

currently under way (Bozkurt et al., 2014).

Chapter III

82

Several disadvantages with these products make the further development of synthetic nerve grafts

with specific biological signals desirable: The AxoGen scaffolds still require harvesting of human

tissue and result in limited control over the material properties. The Matricel scaffold has mostly

the signaling properties of collagen type I via α2β1 and integrins (Bozkurt et al., 2007; Knight et

al., 1999). While the constructs presented here are not likely to be superior to these products, my

approach opens the possibility to use completely synthetic hydrogels and fibers and to endow

them with signaling peptides designed to activate integrin receptors. Further improvements of the

present scaffold may be achieved by additional functionalization of the electrospun fibers

(Bockelmann et al., 2011; Grafahrend et al., 2011) and/or the application of growth factors that

are gradually released (Lee, 2003; Solorio and Zwolinski, 2010)

5 Conclusion

Based on the recovery of motor functions and on physiological assessment of the regenerated

nerves the novel structured implants represent an improvement on conventional hollow conduits

and offer possibilities for further optimization. By elaboration of structural properties (e.g. fiber

density) or by providing more information on the interaction of the involved cell types with

biological signals it should be possible to further narrow the functional gap between artificial

implants and autologous nerve grafts.

Comparison to other studies

83

General Discussion

Peripheral nerve lesions not only may have drastic consequences for the affected individual but, at

the level of the population, also considerable socio-economic impact. This is particularly due to

the high follow up costs generated by these injuries. Reasons for that are, for example, the long

term physical therapy the patient has to undergo to support functional recovery. Furthermore,

until this is achieved to a satisfactory degree, if at all, the patient will not be at his full

professional capacity. To facilitate best possible recovery for the patient and reduce the total costs

for the society, it is necessary to have treatments available with a prospect for recovery at least on

the level of the autologous nerve graft. However, as stated before, a direct end-to-end

neurorrhaphy is not always possible, nor is an autologous nerve graft always available. Hence, at

the moment there are cases were the physician in charge has to make use of alternatives that

might be less promising than the aforementioned solutions, e.g. an artificial implant. Therefore,

there is great interest in improving the performance of artificial implants especially for the

treatment of bigger nerve gaps.

To achieve this, different strategies can be applied, as elaborated in Chapter I of this thesis. The

approach presented here aimed for the integration of sub-micron scale fibers as a guidance

structure into a hollow conduit to obtain a cell-free artificial nerve implant with improved

performance, compared to the hitherto available conventional implants. For this, I designed and

established an electrospinning device for the production of 3D arrays of aligned fibers. These

arrays were embedded in a hydrogel matrix to form a scaffold with an internal guidance structure.

After showing via in vitro experiments that the fibers keep their guidance properties inside the

hydrogel, fiber/gel constructs were incorporated into collagen tubes to obtain implantable devices.

Subsequently these were used to bridge a 15 mm rat sciatic nerve gap in vivo. It showed that, even

though the autologous nerve graft still performed notably better, the incorporation of the fiber/gel

constructs proved to be beneficial for the outcome of regeneration compared to hollow collagen

tubes. Thus I consider this approach a promising basis for further improvements. However, there

have also been other approaches in the past years, of which a representative selection will be

presented briefly in the following.

Comparison to other studies

The suitability of electrospun fibers as guidance structures for peripheral nerve regeneration has

also been demonstrated in other studies. In two consecutive in vivo studies with rats by Ravi V.

Bellamkonda‘s group (Kim et al., 2008; Clements et al., 2009), films of longitudinally aligned

fibers from PAN-MA were incorporated into polysulfone nerve conduits to successfully bridge a

tibial nerve gap of 17 and 14 mm, respectively. In the first study they also performed nerve repair

with an empty (saline-filled) polysulfone conduit and an autologous nerve graft, which, to a

General Discussion

84

certain extent, allows a comparison between their approach and mine. In their study, as in mine,

the best regeneration was found in the group of animals that was treated with an autologous nerve

graft. This was relatively closely followed by the aligned fiber-film group, and the differences

between autograft group and aligned fiber-film group seem smaller than in the present study. This

observation can be explained by several reasons. First of all, the more distal location of the nerve

defect (tibial versus sciatic nerve) generally has a better prediction for functional recovery. In

addition, the outcome of regeneration was assessed only 16 weeks after implantation, which is

four weeks later than in the present study. Surprisingly, in contrast to my study, they did not find

any signs of regeneration in the empty conduits even though the nerve gap to be bridged was only

2 mm longer in their study (17 versus 15 mm). This creates the impression that differences

between autograft group and aligned fiber-film group are smaller than observed in the present

study, allthough this might not be true. However, the data in general suggest a slightly better

performance of their approach compared to mine.

Another example for the successful application of electrospun fibers to perform a PN repair in rats

is the work by Jha et al. (2011). The unique feature of this work is the very convenient procedure

to obtain the internally structured implants. By electrospinning onto two rounded target electrodes

a bundle of aligned fibers was created. In a second step an exterior coating was added to the

construct by electrospraying PGA:PLA copolymer onto the outside of the bundle and thus

forming the final construct. Unfortunately they only used a nerve gap size of 10 mm for biological

testing and stopped the experiment seven weeks after implantation. At this time point the animals

showed a withdrawal reflex in response to foot pinch, indicating sensory reinnervation. This

coincides with the regain of grip strength in the present study, but a more detailed comparison is

hindered by the differences in experimental conditions.

Besides the application of fibers, specifically processed hydrogels have also been used to

implement directional cues into nerve conduits. One way to achieve this was demonstrated by

Yao et al. (2010). They produced collagen nerve conduits with multiple internal channels with

diameters of approximately 500 µm to limit axonal dispersion inside the conduit and thus reduce

inappropriate target reinnervation. While this aspect of the study seemed to be a success, as they

found that even fewer neurons showed double projections to the tibial and the peroneal branch in

the 4-channel group as in the autograft group, the actual regenerative outcome did not differ

between their implants and a conventional hollow collagen conduit when used to repair a 10 mm

sciatic nerve gap in rats.

An alternative approach was used by Bozkurt et al. (2012) who employed a conduit material

called Perimaix. This material, developed by Matricel GmbH (Herzogenrath, Germany), is

derived from porcine collagen that was further processed through cross-linking and directional

freezing to create longitudinal pores for cellular and axonal guidance (Bozkurt et al., 2009).

Bozkurt and coworker used the Perimaix conduit to bridge a 2 cm sciatic nerve gap in rats. This is

Possible augmentation of implant design

85

expected to be a challenging gap size for the application of artificial implants and probably the

largest possible distance to repair in this animal model. Nevertheless they were able to show that

six weeks after implantation axons had crossed the complete implant and depicted the same

diameters as in the autograft group. However, axon density and myelin thickness was reduced to

approximately 50 and 66 %, respectively, of that determined for the autograft group

demonstrating that the nerve transplant was still superior to the artificial scaffold. Functional

recovery was not assessed in this study. A notable difference to most other studies, which might

have had a positive effect on regeneration is the absence of a conventional tubing material.

Instead they opened the nerves epineurium by a longitudinal incision at the site of nerve defect

and later on used it to ensheath the inserted implant, embedding it in a much more natural

environment. Again, the differences in the experimental conditions in addition to the lack of a

comparable negative control (empty tube) do not permit a detailed comparison between this and

the study presented here. However the quality of axonal regeneration across such a defect size is

impressive and led to consecutive clinical testing of the material in humans (Bozkurt et al., 2014).

To conclude, one first of all one has to recognize that comparison between different studies is

difficult in general. Even though the selection presented here is restricted to only one of several

animal species (e.g. rabbit, dog, sheep) used for PN repair studies, there is still a considerable

variability of experimental conditions impeding detailed comparison. These are, for example, the

choice of nerve, defect size, time of recovery and methods for its assessment. Thus, even though

there is a lot of research going on in the field, it is difficult to rank the success of my own

approach in relation to alternative ones. A demand for standardization of peripheral nerve repair

studies with rodents has already been voiced by others (Bozkurt et al., 2012). This would greatly

help to identify the most promising strategies, discard others and accelerate progress in this field

of research.. Apart from that, the selection of studies presented here suggests that my implant

design is not generally superior but competitive to other approaches with internal guidance

structures. However, it seems that the incorporation of physical/topographical guidance cues only

is not sufficient to achieve results on the level of autologous nerve grafts. Further augmentation of

functionality seems to be necessary to obtain a true alternative to the current “gold standard“, the

autologous nerve graft.


The arrangement of guidance cues can influence an implant’s performance in terms of

regeneration as shown, for example, by Clements et al. (2009). Thus a further improvement of my

implant’s performance by adjustment of certain parameters is conceivable, e.g. by increasing the

amount of fibers and their dispersion inside the implant. However, in the past decade there has

been a variety of different design approaches for artificial nerve conduits with or without

General Discussion

86

incorporated physical/topographical guidance cues, and, to my knowledge, none of them was able

to compete with the autologous nerve graft with respect to regeneration and functional recovery in

gap sizes bigger than 1 cm. Apparently the incorporation of physical/topographical guidance cues

has beneficial effect, but further augmentation, e.g. through implementation of biological

signaling, seems necessary to obtain results equivalent to what can be achieved with autologous

nerve grafts.

There are several possibilities to equip an implant with biological signals that support axonal

regeneration. An obvious way to achieve this is preseeding the conduit with appropriate cells

before implantation. The positive effect of preseeded cells was shown in numerous studies with a

variety of cell types (di Summa et al., 2010; Ding et al., 2010; Bozkurt et al., 2012; Allodi et al.,

2014; Beigi et al., 2014). As an example, by preseeding the Perimaix conduits with Schwann cells

for repairing a 2 cm sciatic nerve lesion in rats Bozkurt et al. (2012) achieved regeneration that

was almost as effective as when an autologous nerve graft was used. However, for the application

in humans, which is the final therapeutic goal of all this research, preseeded conduits involve

certain complications, as discussed earlier (Chapter I 3.4). Nevertheless, it is conceivable to

mimic certain features those cells exhibit to support regeneration.

One of those features is the secretion of ECM molecules. The ECM is equipped with a variety of

molecular components that can be recognized by axonal receptors, e.g. integrins. Hence, a way to

improve the performance of artificial implants is the incorporation of selected ECM components.

In the present study, I used a rather simple approach to achieve an enhancement of the fibers

guidance properties, namely by blending the PCL solution with collagen, a major ECM

constituent. This was shown to have positive effects on Schwann cell migration and axonal

regeneration in vitro (Schnell et al., 2007; Mey et al., 2012) and also led to slightly better

performance in my in vitro experiments compared to pure PCL. More sophisticated methods

involve chemical modification of materials by covalent binding of signaling peptides, e.g. an

RGD sequence (arginine-glycine-aspartic acid), which cells can bind to via certain integrins (Koh

et al., 2008; Bockelmann et al., 2011; Grafahrend et al., 2011). However, an advantage of

covalent binding of signaling peptides to the guidance structure over mere blending of petides to

the electrospinning solution could not be shown yet.

Another beneficial feature cells exhibit to support regeneration is the release of growth factors.

There is a variety of different growth factors known to support the regeneration of axons after

PNI, e.g. NGF, GDNF, NT-3, CNTF etc. (Kokai et al., 2010, 2011; Allodi et al., 2011).

Depending on the factor, the positive effect on regeneration is mediated through different

mechanisms, for example the promotion of axonal outgrowth or neuronal survival (de Ruiter et

al., 2009). Thus, by the combination of different growth factors, synergistic effects are to be

expected that could lead to better support than the application of only a single type of growth

factor (Madduri et al., 2010a). However more comprehensive analysis of the dose-dependency of


87

different factors and their possible interactions are desirable to choose the best combination of

growth factors for an artificial nerve conduit. Possibilities for the incorporation of growth factors

into artificial conduits have already been presented in a number of studies. Usually systems for

the sustained delivery of factors over a longer period of time are preferred. This can be achieved

by encapsulating the factors into slowly degradable materials, e.g. PCL or cross-linked gelatin

(Xu et al., 2003; Moore et al., 2006; Chang, 2009; Valmikinathan et al., 2009; Solorio and

Zwolinski, 2010; Kokai et al., 2011; Liu et al., 2011).

Conclusion

As shown by the experiments presented in this thesis, I successfully designed and manufactured a

cell-free artificial nerve conduit for the support of peripheral nerve regeneration. The new design

with the incorporated fiber/gel scaffold represents an improvement over a hollow collagen

conduit. As the performance of my implant design with respect to the support of regeneration and

functional recovery is generally on a comparable level with other designs presented so far, further

effort in advancing the design for the use in humans is justified. To reach this goal I suggest the

following steps to be taken.

First of all, the physical structure of the internal scaffold should be optimized. The distribution of

fibers inside the implants, used in the present study, was not homogenous. The fibers were mainly

localized in the center of the scaffold in an area of approximately 1 mm width and 0.5 mm heights

(when looking at a cross section). Here, one should strive for an adjustment of the procedure for

fiber embedding and incorporation into the tube to obtain a more homogenous distribution of

fibers throughout the whole cross section of the conduit. In my opinion a moderate increase in

fiber density would also be favorable. This might prove beneficial for regeneration, while

detrimental effects are not to be expected. At least I conclude this from comparisons with other

studies were the process for manufacturing of the implants suggests a final fiber density that was

higher than in the case presented here, e.g. (Jha et al., 2011).

Secondly, biological signaling should be implemented into the conduit. To obtain the best

possible results, I suggest to use a combination of functionalizing the fibers through modification

of their surface with ECM molecules, as suggested by (Grafahrend et al., 2011), and the

incorporation of a release system for growth factors. For this purpose, the use of microspheres

encapsulating the growth factors seems reasonable, since they are already widely used in this

context and can be incorporated into the conduit wall (Kokai et al., 2010). Additionally I suggest

the use of cocktail of different growth factor to make use of synergistic effects. However ideal

combinations and concentrations remain to be determined.

Finally, before testing the optimized implants in humans, a second animal study in vivo, using the

optimized implants, would be desirable. In this case, I suggest to choose a more challenging gap

Conclusion

88

size to be repaired. This way one should obtain more distinct differences between experimental

groups, which eases the evaluation of the implants performance. However this requires a bigger

animal model for example, e.g. Beagle dogs (Wang et al., 2005; Ding et al., 2010; Huang et al.,

2013). Also, in contrast to the present study, the assessment of functional recovery should be

expanded to sensory function to give a more detailed insight into different aspects of regeneration.

A successful completion of the suggested measures would, in my opinion, justify the start of

clinical trials in humans.

89

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Abbreviations

BDNF Brain-derived neurotrophic factor

CAM Cell adhesion molecule

CNS Central nervous system

CNTF Ciliary neurotrophic factor

DRG Dorsal root ganglion

ECM Extracellular matrix

ePTFE Extended poly(tetrafluoroethylene)

ErbB Erythroblastic leukemia viral oncogene (epidermal growth factor receptor

family)

FGF Fibroblast growth factor

GDNF Glial-derived neurotrophic factor

GFR Receptor of GDNF-family ligands

GGF Glial growth factor

IGF Insulin-like growth factor

IL Interleukin

LIF Leukemia inhibitory factor

NGF Nerve growth factor

NT Neurotrophin

PAN-MA Poly(acrylonitrile-co-methacrylate)

PCL Poly(ε-caprolactone)

PCLEEP Poly(caprolactone-co-ethyl ethylene phosphate)

PGA Poly(glycolic acid)

PHB Poly(hydroxybutyric acid)

pHEMA Poly(2-hydroxyethyl methacrylate)

pHEMA-MMA Poly(2-hydroxyethyl methacrylate-co-methyl methacrylate)

PLA Poly(lactic acid); the D- and L-isomers are often referred to as PDLA,

PLLA, respectively

PLGA Poly(lactic-co-glycolic acid)

PN Peripheral nerve

PNI Peripheral nerve injury

PNS Peripheral nervous system

PPE Polyphosphoester

PPy Poly(pyrrole)

sPEG Star-shaped poly(ethylene glycol)

STAT Signal transducer and activator of transcription

TGF Transforming growth factor

trk Tyrosine kinase/tropomyosin receptor kinase (neurotrophin receptor)

VEGF Vascular endothelial growth factor

109

Figures and Tables

Figure 1 Peripheral nerve degeneration and regeneration. ..................................................... 11

Figure 2 Design strategies for artificial nerve bridges. ............................................................. 18

Figure 3 Examples of artificial nerve guides that were successfully tested in animal

experiments. .................................................................................................................................. 20

Figure 4 Electrospinning of aligned nanofibers. ....................................................................... 22

Figure 5 Evaluation of artificial nerve implants in animal experiments. ................................ 33

Figure 6 Functionalization of polymer fibers for nerve regeneration. .................................... 36

Figure 7 Conduit design for artificial nerve guides. .................................................................. 42

Figure 8 Electrospinning setup for the collection of 3D fiber arrays....................................... 43

Figure 9 Three-dimensional arrays of parallel fibers. .............................................................. 44

Quantitative characterization of the 3D fiber array. .................................................................. 48

Figure 11 Embedding of 3D fiber stacks in collagen. ................................................................ 49

Figure 12 Gel-embedded PCL fibers affect cell morphology. ................................................... 50

Figure 13 Gel-embedded PCL fibers as a guidance structure for Schwann cells and sensory

neurons. ......................................................................................................................................... 51

Figure 14 Quantification of the guidance properties of gel-embedded PCL fibres. ............... 52

Figure 15 Integration of the composit scaffold into a biocompatible conduit. ....................... 54

Figure 16 Brief documentation of the fiber production and embedding in gelatin. .............. 61

Figure 17 Tissue regeneration following sciatic nerve transection (15 mm gap size) and

subsequent repair using different nerve bridges. ...................................................................... 65

Figure 18 Cross sectional area of the regenerated tissue in the medial segment 12W after

nerve transection and repair. ...................................................................................................... 66

Figure 19 Schwann cell infiltration and axonal regeneration at the proximal interface

between nerve and implant 2W post surgery. ........................................................................... 67

Figure 20 Schwann cell infiltration and axonal regeneration in the medial segment of the

regenerated nerve 12W post surgery. ........................................................................................ 68

Figure 21 Quantitative analysis of the regenerated nerve tissue. ........................................... 69

Figure 22 Similar percentage of regenerating motoraxons in all implants. ........................... 70

Figure 23 Electromyography performed 12W after sciatic nerve repair. .............................. 73

Figure 24 Behavioral analysis of nerve regeneration. .............................................................. 76

Figure 25 Determination of muscular atrophy. ........................................................................ 77

Table 1 Compound muscle action potentials after PN repair with artificial implants ........... 74

Table 2 Muscle weight normalized to the corresponding muscle on the contralateral side 2

and 12 weeks after respective treatment. .................................................................................. 78

111

Prior Publications and Contributions

The experiments reported in this thesis represent a joint effort of myself and several scientists,

who contributed to the investigation, supervised me or were supervised by me. I performed the

majority of all experimental work in the laboratory and I had a key role in designing, evaluating

and writing the studies that have already been submitted for publication. This is documented by

the respective lists of co-authors (see below).

Animal surgery was carried out by Gary Brook (with my assistance), programming and design of

the electrophysiology was the work of Thomas Kuenzel.

Electrospun Fibers as Substrates for Peripheral Nerve Regeneration

Jörg Mey, Gary Brook, Dorothee Hodde and Andreas Kriebel

Published in Advanced Polymer Science (2012) 246: 131–170

Three-dimensional configuration of orientated fibers as guidance structures for cell

migration and axonal growth

Andreas Kriebel, Muhammad Rumman, Miriam Scheld, Dorothee Hodde, Gary Brook and Jörg

Mey

Published in Journal of Biomedical Materials Research: Part B (2014) 102B: 356-365

Cell-free artificial nerve implants with electrospun fibers in a three-dimensional gelatin

matrix support nerve regeneration in vivo

Andreas Kriebel, Dorothee Hodde, Thomas Kuenzel, Jessica Engels, Gary Brook and Jörg Mey

Submitted for publication (March 2015)

113

Acknowledgments

Zuallererst möchte ich mich bei Jörg bedanken, dass er mir das Vertrauen geschenkt hat, dieses

Projekt zu bearbeiten. Ohne immer einer Meinung zu sein, sind wir dennoch gut miteinander

ausgekommen. Es hat Spaß gemacht mit Dir zu arbeiten.

Bedanken möchte ich mich auch bei Hermann, der mir uneigennützig Labor und Arbeitsplatz zur

Verfügung gestellt hat und auch sonst stets ein offenes Ohr für mich hatte.

Two persons that contributed essentially to this project are Gary and Doro. I dont know, whether I

could have spend all this time working with the animals without having Doro to talk to. I could

always rely on her and she helped whenever she could. The same accounts to Gary. I had a great

time assisting him with the operations. I have rarely seen a PI contributing so much to the

practical parts of a project. Besides that, he kept up my motivation, when things did not work out

as supposed to. That really helped a lot. You were a great team to work with, I hope I was able to

return something to you.

Thomas und Marcus danke ich für die vielen Ratschläge, die ich mir bei Ihnen geholt habe. Vor

allem Thomas tatkräftige Unterstützung hat nennenswert zum erfolgreichen Abschluss dieser

Arbeit beigetragen.

Darüber hinaus möchte ich auch den Bachelorstudenten danken, die durch ihre Abschlussarbeiten

Anteil an diesem Projekt hatten.

Natürlich darf auch der Rest des Instituts nicht vergessen werden. Ich habe mich in den nunmehr

fast sechs Jahren, die ich hier tätig war immer sehr wohl gefühlt. Besonders bei Michael muss ich

mich bedanken. Er hat zwar nie Zeit, aber nimmt sie sich doch immer, wenn man seine Hilfe

braucht, und die brauchte ich öfters. Roland muss ich auch danken, denn ohne seine Kompetenz

im Umgang mit Word wären vermutlich einige meiner PC-Komponenten aus dem Fenster

geflogen. Philipp, Lutz und Julius muss ich nicht danken…die wissen schon, dass sie für mich

mehr als bloß Kollegen sind und ich hoffe, dass es auch noch lange so bleibt.

Bleibt noch der Dank an meine Familie.

Katharina, Maya und Adrian…ihr seid der Antrieb und gleichzeitig die willkommenste

Ablenkung in meinem Leben. Meinem Vater danke ich dafür, dass er mir immer meine

Möglichkeiten aufgezeigt hat, ohne meine Entscheidungen treffen zu wollen.

Zu guter Letzt bedanke ich mich bei meiner Mutter. Sie hat mich zu dem Menschen gemacht hat,

der ich heute bin. Ich habe nun wohl fast alles erreicht, das sie sich für mich gewünscht hat. Ich

wünschte, sie hätte es erleben können.

Dieses Projekt wurde von der Deuschen Forschungsgemeinschaft finanziert (grant ME 1261/11-1)

115

Curriculum Vitae

Persönliche Daten

Name Andreas Kriebel

Geburtstag 11.12.1981

Geburtsort Leverkusen

Staatsangehörigkeit Deutsch

Werdegang

2001 Erlangung der allgemeinen Hochschulreife an der Landrat-

Lucas-Schule in Leverkusen-Opladen

2001-2002 Grundwehrdienst im Fallschirmjägerbataillon 263 in

Zweibrücken

2002-2003 Studium des Maschienenbaus, RWTH Aachen University

2003-2010 Studium der Biologie, RWTH Aachen University

Abschluss: Diplom-Biologe

2010-2014 Wissenschaftlicher Angestellter

Institut für Biologie II, RWTH Aachen University

2010-2015 Promotionsstudium, RWTH Aachen University

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